Biomimmetic nanofiber scaffold for soft tissue and soft tissue-to-bone repair, augmentation and replacement

ABSTRACT

An implantable device is provided for soft-tissue or soft tissue-to-bone repair, fixation, augmentation, or replacement that includes a biomimetic and biodegradable nanofiber scaffold. Also provided is a fully synthetic implantable multiphased scaffold which includes, in a single continuous construct, a plurality of phases to mimic the natural anatomy of a tendon or ligament and their insertion sites. Also provided are scaffold apparatuses for musculoskeletal tissue engineering.

This application is a divisional of U.S. Ser. No. 12/583,072, filed Aug.12, 2009, now U.S. Pat. No. 8,753,391 which is a continuation-in-part ofand claims the benefit of (a) PCT International Application No.PCT/US2008/001889, filed Feb. 12, 2008 which in turn claimed priority ofU.S. Provisional Applications Nos. 60/901,047 and 60/905,649, filed Feb.12, 2007 and Mar. 7, 2007, respectively; (b) PCT InternationalApplication No. PCT/US2008/007323, filed Jun. 11, 2008 which in turnclaimed priority of U.S. Provisional Application No. 60/934,198, filedJun. 11, 2007 and (c) PCT International Application No.PCT/US2008/007357, filed Jun. 11, 2008 which in turn claimed priority ofU.S. Provisional Application No. 60/934,182, filed Jun. 11, 2007. Thisapplication also claims the benefit of (d) U.S. Provisional ApplicationNo. 61/215,085, filed May 1, 2009. The entire contents of each of (a)through (d) are hereby incorporated by reference herein.

This invention was made with government support under NSF GraduateFellowship (GK-12 0338329) awarded by the National Science Foundationand Grant Numbers R21 AR052402-01A1, R01 AR056459-01 and R21 AR055280-01awarded by the National Institutes of Health. The government may havecertain rights in the invention.

BACKGROUND

Throughout this application, certain publications are referenced. Fullcitations for these publications, as well as additional relatedreferences, may be found immediately preceding the claims. Thedisclosures of these publications are hereby incorporated by referenceinto this application in order to more fully describe the state of theart as of the date of the invention described and claimed herein.

This application relates to musculoskeletal tissue engineering. Forexample, a scaffold apparatus is discussed below which can serve as afunctional interface between multiple tissue types. Methods forpreparing a multi-phase scaffold are also discussed. Some exemplaryembodiments which include a soft tissue-bone interface are discussed.

As examples of soft tissue-bone interface, the human rotator cuff andanterior cruciate ligament (ACL) are described below. The rotator cuff,the ACL and the ACL-bone interface are used in the following discussionas examples to aid in understanding the description of the methods andapparatuses of this application. This discussion, however, is notintended to, and should not be construed to, limit the claims of thisapplication.

The Rotator Cuff

The rotator cuff consists of a group of four muscles and tendons,including the supraspinatus, infraspinatus, teres minor, andsubscapularis, which function in synchrony to stabilize the glenohumeraljoint as well as to actively control shoulder kinematics. Thesupraspinatus tendon inserts into the humeral head via a directinsertion exhibiting region-dependent matrix heterogeneity and mineralcontent.

Four distinct yet continuous tissue regions are observed at thetendon-bone junction (FIG. 9A): tendon proper, non-mineralizedfibrocartilage, mineralized fibrocartilage and bone (Benjamin, 1986;Benjamin, 2002; Woo, 1988). The tendon proper consists of fibroblastsfound between aligned collagen fibers in a matrix rich in collagen I,with small amounts of collagen III and proteoglycans (Blevins, 1997).The non-mineralized fibrocartilage region is composed offibrochondrocytes in a matrix of collagen I, II, and III with fibersoriented perpendicular to the calcified interface region (Kumagai,1994). The mineralized fibrocartilage region consists of hypertrophicfibrochondrocytes within a matrix of collagen I and II (Kumagai, 1994)as well as collagen X (Thomopoulos, 2003). The last region of theinsertion site is bone which consists of osteoblasts, osteoclasts, andosteocytes in a mineralized matrix rich in type I collagen.

This controlled matrix heterogeneity exhibited by the tendon-boneinterface serves to minimize stress concentrations and to mediate loadtransfer between two distinct tissue types (Thomopoulos, 2003; Woo,1988). Due to its functional significance, interface regeneration is apre-requisite for biological fixation.

Rotator cuff tears are among the most common injuries afflicting theshoulder, with greater than 75,000 repair procedures performed annuallyin the United States alone (Vitale, 2007). Clinical intervention isrequired because injuries to the rotator cuff do not heal, largely dueto the complex anatomy noted above and the extended range of motion ofthe shoulder joint, as well as the relative weakening andhypovascularization of the cuff tendons (Codman, 1934; Yamanaka, 1994;Dejardin, 2001). Moreover, chronic degeneration increases both thefrequency and size of cuff tears with age (Tempelhof, 1999) and isconsidered the main contributing factor in the pathogenesis of rotatorcuff tendon tears (Dejardin, 2001; Soslowsky, 2000). Early primaryanatomic repair followed by carefully controlled rehabilitation iscurrently the treatment of choice for symptomatic rotator cuff tears(Dejardin, 2001).

Rotator cuff repair has evolved from traditional open repair to“mini-open” to primarily arthroscopic (Gartsman, 2001; Galatz, 2004;Willaims, 2004; Mazzocca, 2005; Cole 2007). This transition has occurreddue to advances in surgical techniques and fixation methods, with thecurrent technique being a double row “suture-bridge” technique whichsimulates the compression afforded by transosseous sutures previouslyused in open and mini-open repairs (Park, 2007). These methods have beenshown to improve mechanical strength and graft stability (Park, 2005).The focus in the field now centers on how to address the challenge ofachieving functional rotator cuff healing and/or augmentation, which isessential for long term clinical success.

Currently, a significant demand exists for a functional tendon graftingsystem which can augment and promote rotator cuff healing post surgicalrepair, due to the relatively high failure rates associated with currentrepair procedures as well as the clinical need to treat large tears andchronic degeneration of the rotator cuff tendons. For example, failurerates as high as 90% have been reported after primary repair of chronicrotator cuff injuries (Galatz, 2004), generally attributed to factorssuch as osteoporotic bone, degenerative and poorly vascularized tendons,severe tendon weakening, muscle atrophy, and size of the original defect(Gazielly, 1994; Gartsman, 1997; Mansat, 1997; Rokito, 1996; Romeo,1999). Moreover, the primary repair of chronic degenerative cuffinjuries often results in excessive tension on the cuff tissues and atthe repair site (Dejardin, 2001; DeOrio, 1984; Gerber, 1999). To improvehealing, synthetic grafts (Post, 1985; Ozaki, 1986) have been designedto reconstruct large rotator cuff defects. See also, e.g., U.S. Pat. No.7,112,417. However, these devices are suboptimal due to concerns ofbiocompatibility as well as their inability to meet the functionaldemand of the native tendon.

Recently, biological matrices such as acellularized allogeneic andxenogeneic extracellular matrix scaffolds have emerged as promisinggrafts for rotator cuff repair and augmentation (Dejardin, 2001;Schlpegel, 2006). Both collagen-rich dermis and small intestinalsubmucosa (SIS) (Badylak, 2002) have been marketed commercially as graftpatches for reinforcing soft tissue repair following rotator cuffsurgery (Derwin, 2006; Lannotti, 2006). See also, e.g., U.S. Pat. Nos.6,638,312 and 7,160,333. SIS is particularly attractive as it exhibits abiomimetic, collagen nanofiber-based architecture and alignment, thus itcan be readily remodeled by host cells while encouraging angiogenesisand neo-collagen production (Badylak, 2002).

Highly promising results have been reported for SIS in several animalmodels (Dejardin, 2001; Schlegel, 2006), but unfortunately suboptimaloutcomes were observed in human trials (Lannotti, 2006; Sclamberg,2004), in which augmentation with SIS did not improve the rate of tendonhealing or clinical outcome scores. Similar outcomes have been reportedfor other biological grafts used in rotator cuff repair (Sclamberg,2004; Coons, 2006).

The suboptimal results of biologically-derived grafts may be attributedto mismatch in mechanical properties and the rapid matrix remodelingexperienced in the physiologically demanding and often diseased shoulderjoint. Utilizing a canine model, Derwin et al. (Derwin, 2006) performeda systematic comparison of the biomechanical properties of commerciallyavailable extracellular matrices for rotator cuff augmentation. Amismatch in mechanical properties with the canine infraspinatus tendonwas observed for all types of extracellular matrix tested.

Moreover, it has been reported that the mechanical properties of SISdecreased as resorption occurred prematurely (Derwin, 2006). Thus themismatch in the kinetics of graft remodeling and neo-collagen formationcompromised the clinical outcome. Therefore, the debilitating effect ofrotator cuff tears coupled with the high incidence of failure associatedwith existing graft choices emphasize the clinical need for functionalrotator graft augmentation solutions.

The Anterior Cruciate Ligament (ACL)

The ACL consists of a band of regularly oriented, dense connectivetissue that spans the junction between the femur and tibia. Itparticipates in knee motion control and acts as a joint stabilizer,serving as the primary restraint to anterior tibial translation. Thenatural ACL-bone interface consists of three regions: ligament,fibrocartilage (non-mineralized and mineralized) and bone. The naturalligament to bone interface is arranged linearly from ligament tofibrocartilage and to bone. The transition results in varying cellular,chemical, and mechanical properties across the interface, and acts tominimize stress concentrations from soft tissue to bone.

The ACL is the most often injured ligament of the knee. Due to itsinherently poor healing potential and limited vascularization, ACLruptures do not heal effectively upon injury, and surgical interventionis typically needed to restore normal function to the knee. Clinically,autogenous grafts based on either bone-patellar tendon-bone (BPTB) orhamstring-tendon (HST) grafts are often a preferred grafting system forACL reconstruction, primarily due to a lack of alternative graftingsolutions. Current ACL grafts are limited by donor site morbidity,tendonitis and arthritis. Synthetic grafts may exhibit good short termresults but encounter clinical failure in long-term follow-ups, sincethey are unable to duplicate the mechanical strength and structuralproperties of human ACL tissue. ACL tears and ruptures are currentlycommonly repaired using semitendinosus grafts. Although semitendinosusautografts are superior, they often fail at the insertion site betweenthe graft and the bone tunnel. One of the major causes of failure inthis type of reconstruction grafts is its inability to regenerate thesoft-tissue to bone interface.

Despite their distinct advantages over synthetic substitutes, autogenousgrafts have a relatively high failure rate. A primary cause for the highfailure rate is the lack of consistent graft integration with thesubchondral bone within bone tunnels. The site of graft contact infemoral or tibial tunnels represents the weakest point mechanically inthe early post-operative healing period. Therefore, success of ACLreconstructive surgery depends heavily on the extent of graftintegration with bone.

ACL reconstruction based on autografts often results in loss offunctional strength from an initial implantation time, followed by agradual increase in strength that does not typically reach the originalmagnitude. Despite its clinical success, long term performance ofautogenous ligament substitutes is dependent on a variety of factors,including structural and material properties of the graft, initial grafttension, intrarticular position of the graft, as well as fixation of thegraft. These grafts typically do not achieve normal restoration of ACLmorphology and knee stability.

There is often a lack of graft integration with host tissue, inparticular at bony tunnels, which contributes to suboptimal clinicaloutcome of these grafts. The fixation sites at the tibial and femoraltunnels, instead of the isolated strength of the graft material, havebeen identified as mechanically weak points in the reconstructed ACL.Poor graft integration may lead to enlargement of the bone tunnels, andin turn may compromise the long term stability of the graft.

Increased emphasis has been placed on graft fixation, as post surgeryrehabilitation protocols require the immediate ability to exercise fullrange of motion, reestablish neuromuscular function and weight bearing.During ACL reconstruction, the bone-patellar tendon-bone orhamstring-tendon graft is fixed into the tibial and femoral tunnelsusing a variety of fixation techniques. Fixation devices include, forexample, staples, screw and washer, press fit EndoButton® devices, andinterference screws. In many instances, EndoButton® devices or Mitek®Anchor devices are utilized for fixation of femoral insertions. Staples,interference screws, or interference screws combined with washers can beused to fix the graft to the tibial region.

Recently, interference screws have emerged as a standard device forgraft fixation. The interference screw, about 9 mm in diameter and atleast 20 mm in length, is used routinely to secure tendon to bone andbone to bone in ligament reconstruction. Surgically, the knee is flexedand the screw is inserted from the para-patellar incision into thetibial socket, and the tibial screw is inserted just underneath thejoint surface. After tension is applied to the femoral graft and theknee is fully flexed, the femoral tunnel screw is inserted. Thisprocedure has been reported to result in stiffness and fixation strengthlevels which are adequate for daily activities and progressiverehabilitation programs.

While the use of interference screws have improved the fixation of ACLgrafts, mechanical considerations and biomaterial-related issuesassociated with existing screw systems have limited the long termfunctionality of the ligament substitutes. Screw-related laceration ofeither the ligament substitute or bone plug suture has been reported. Insome cases, tibial screw removal was necessary to reduce the painsuffered by the patient. Stress relaxation, distortion of magneticresonance imaging, and corrosion of metallic screws have providedmotivation for development of biodegradable screws based onpoly-α-hydroxy acids. While lower incidence of graft laceration wasreported for biodegradable screws, the highest interference fixationstrength of the grafts to bone is reported to be 475 N, which issignificantly lower than the attachment strength of ACL to bone. Whentendon-to-bone fixation with polylactic acid-based interference screwswas examined in a sheep model, intraligamentous failure was reported by6 weeks. In addition, fixation strength is dependent on quality of bone(mineral density) and bone compression.

Two insertion zones can be found in the ACL, one at the femoral end andanother located at the tibial attachment site. The ACL can attach tomineralized tissue through insertion of collagen fibrils, and thereexists a gradual transition from soft tissue to bone. The femoralattachment area in the human ACL was measured to be 113±27 mm² and136±33 mm² for the tibia insertion. With the exception of the mode ofcollagen insertion into the subchondral bone, the transition from ACL tobone is histologically similar for the femoral and tibial insertionsites.

The insertion site is comprised of four different zones: ligament,non-mineralized fibrocartilage, mineralized fibrocartilage, and bone.The first zone, which is the ligament proper, is composed of solitary,spindle-shaped fibroblasts aligned in rows, and embedded in parallelcollagen fibril bundles of 70-150 μm in diameter. Primarily type Icollagen makes up the extracellular matrix, and type III collagen, whichare small reticular fibers, are located between the collagen I fibrilbundles. The second zone, which is fibro-cartilaginous in nature, iscomposed of ovoid-shaped chondrocyte-like cells. The cells do not liesolitarily, but are aligned in rows of 3-15 cells per row. Collagenfibril bundles are not strictly parallel and much larger than thosefound in zone 1. Type II collagen is now found within the pericellularmatrix of the chondrocytes, with the matrix still made up predominantlyof type I collagen. This zone is primarily avascular, and the primarysulfated proteoglycan is aggrecan. The next zone is mineralizedfibrocartilage. In this zone, chondrocytes appear more circular andhypertrophic, surrounded by larger pericellular matrix distal from theACL. Type X collagen, a specific marker for hypertrophic chondrocytesand subsequent mineralization, is detected and found only within thiszone. The interface between mineralized fibrocartilage and subjacentbone is characterized by deep inter-digitations. Increasing number ofdeep inter-digitations is positively correlated to increased resistanceto shear and tensile forces during development of rabbit ligamentinsertions. The last zone is the subchondral bone and the cells presentare osteoblasts, osteocytes and osteoclasts. The predominant collagen istype I and fibrocartilage-specific markers such as type II collagen areno longer present.

For bone-patellar tendon-bone grafts, bone-to-bone integration with theaid of interference screws is the primary mechanism facilitating graftfixation. Several groups have examined the process of tendon-to-bonehealing.

Blickenstaff et al. (1997) evaluated the histological and biomechanicalchanges during the healing of a semitendinosus autograft for ACLreconstruction in a rabbit model. Graft integration occurred by theformation of an indirect tendon insertion to bone at 26 weeks. However,large differences in graft strength and stiffness remained between thenormal semi-tendinosus tendon and anterior cruciate ligament after 52weeks of implantation.

In a similar model, Grana et al. (1994) reported that graft integrationwithin the bone tunnel occurs by an intertwining of graft and connectivetissue and anchoring of connective tissue to bone by collagenous fibersand bone formation in the tunnels. The collagenous fibers have theappearance of Sharpey's fibers seen in an indirect tendon insertion.

Rodeo et al. (1993) examined tendon-to-bone healing in a canine model bytransplanting digital extensor tendon into a bone tunnel within theproximal tibial metaphysis. A layer of cellular, fibrous tissue wasfound between the tendon and bone, and this fibrous layer matured andreorganized during the healing process. As the tendon integrated withbone through Sharpey's fibers, the strength of the interface increasedbetween the second and the twelfth week after surgery. The progressiveincrease in strength was correlated with the degree of bone in growth,mineralization, and maturation of the healing tissue.

In most cases, tendon-to-bone healing with and without interferencefixation does not result in the complete re-establishment of the normaltransition zones of the native ACL-bone insertions. This inability tofully reproduce these structurally and functionally different regions atthe junction between graft and bone is detrimental to the ability of thegraft to transmit mechanical stress across the graft proper and leads tosites of stress concentration at the junction between soft tissue andbone.

Zonal variations from soft to hard tissue at the interface facilitate agradual change in stiffness and can prevent build up of stressconcentrations at the attachment sites.

The insertion zone is dominated by non-mineralized and mineralizedfibrocartilage, which are tissues adept at transmitting compressiveloads. Mechanical factors may be responsible for the development andmaintenance of the fibrocartilagenous zone found at many of theinterfaces between soft tissue and bone. The fibrocartilage zone withits expected gradual increase in stiffness appears less prone tofailure.

Benjamin et al. (1991) suggested that the amount of calcified tissue inthe insertion may be positively correlated to the force transmittedacross the calcified zone.

Using simple histomorphometry techniques, Gao et al. determined that thethickness of the calcified fibrocartilage zone was 0.22±0.7 mm and thatthis was not statistically different from the tibial insertion zone.While the ligament proper is primarily subjected to tensile andtorsional loads, the load profile and stress distribution at theinsertion zone is more complex.

Matyas et al. (1995) combined histomorphometry with a finite elementmodel (FEM) to correlate tissue phenotype with stress state at themedial collateral ligament (MCL) femoral insertion zone. The FEM modelpredicted that when the MCL is under tension, the MCL midsubstance issubjected to tension and the highest principal compressive stress isfound at the interface between ligament and bone.

Calcium phosphates have been shown to modulate cell morphology,proliferation and differentiation. Calcium ions can serve as a substratefor Ca²⁺-binding proteins, and modulate the function of cytoskeletonproteins involved in cell shape maintenance.

Gregiore et al. (1987) examined human gingival fibroblasts andosteoblasts and reported that these cells underwent changes inmorphology, cellular activity, and proliferation as a function ofhydroxyapatite particle sizes. Culture distribution varied from ahomogenous confluent monolayer to dense, asymmetric, and multi-layers asparticle size varied from less than 5 μm to greater than 50 μm, andproliferation changes correlated with hydroxyapatite particles size.

Cheung et al. (1985) further observed that fibroblast mitosis isstimulated with various types of calcium-containing complexes in aconcentration-dependent fashion.

Chondrocytes are also dependent on both calcium and phosphates for theirfunction and matrix mineralization. Wuthier et al. (1993) reported thatmatrix vesicles in fibrocartilage consist of calcium-acidicphospholipids-phosphate complex, which are formed from actively acquiredcalcium ions and an elevated cytosolic phosphate concentration.

Phosphate ions have been reported to enhance matrix mineralizationwithout regulation of protein production or cell proliferation, likelybecause phosphate concentration is often the limiting step inmineralization. It has been demonstrated that human foreskin fibroblastswhen grown in micromass cultures and under the stimulation of lacticacid can dedifferentiate into chondrocytes and produce type II collagen.

Cheung et al. (1985) found a direct relationship betweenβ-glycerophosphate concentrations and mineralization by both osteoblastsand fibroblasts. Increased mineralization by ligament fibroblasts isobserved with increasing concentration of β-glycerophosphate, a mediaadditive commonly used in osteoblast cultures. These reports stronglysuggest the plasticity of the fibroblast response and that thede-differentiation of ligament fibroblasts is a function of mineralcontent in vitro.

Progressing through the four different zones which make up the nativeACL insertion zone, several cell types are identified: ligamentfibroblasts, chondrocytes, hypertrophic chondrocytes and osteoblasts,osteoclasts, and osteocytes. The development of in vitro multi-cell typeculture systems facilitates the formation of the transition zones.

No reported studies on either the co-culture of ligament fibroblastswith osteoblasts, nor on the in vitro and in vivo regeneration of thebone-ligament interface are known.

No reported studies which examine the potential of multi-phasedscaffolds in facilitating the fixation of ligament or tendon to bone areknown. As the interface between graft and bone is the weakest pointduring the initial healing period, recent research efforts in ACL tissueengineering have concentrated on design of multi-phased scaffolds inorder to promote graft integration.

Goulet et al. (2000) developed a bio-engineered ligament model, whereACL fibroblasts were added to the structure and bone plugs were used toanchor the bioengineered tissue. Fibroblasts isolated from human ACLwere grown on bovine type I collagen, and the bony plugs were used topromote the anchoring of the implant within the bone tunnels.

Cooper et al. (2000) and Lu et al. (2001) developed a tissue engineeredACL scaffold using biodegradable polymer fibers braided into a 3-Dscaffold. This scaffold has been shown to promote the attachment andgrowth of rabbit ACL cells in vitro and in vivo. However, no multiphasedscaffolds for human ligament-to-bone interface are known.

SUMMARY

This application describes scaffold apparatuses for musculoskeletaltissue engineering.

While the mechanism for interface regeneration is not known, knowledgeof the structure-function relationship at the tendon-bone insertion(Thomopoulos, 2003; Thomopoulos, 2006) provides invaluable clues inbiomimetic nanofiber scaffold design for interface regeneration.Combining biomechanical testing with the quasi-linear viscoelastic model(QLV) (Fung, 1972), Thomopoulos et al. (Thomopoulos, 2003) determinedthe mechanical properties of the rat supraspinatus tendon insertionsites and later related it to collagen orientation using a finiteelement model (Thomopoulos, 2006). It was found that controlled collagenorganization plays an important role in reducing stress concentration atthe tendon-bone insertion (Thomopoulos, 2006). Specifically, the averagecollagen fiber angle varied from 83-98° in the non-mineralized and86-103° in the mineralized fibrocartilage region, indicating that theinterface fiber architecture deviated minimally from the tendon proper.

In addition to collagen alignment, another intrinsic parameter of theinterface is the region-dependent mineral distribution across theinsertion site (Benjamin, 1986; Woo, 1988). Calcium phosphate is a primemodulator of both the biochemical milieu and the nature of mechanicalstimuli presented to cells. Moreover, the spatial variation in mineralcontent at the interface is mechanically relevant, as increased mineralcontent has been associated with higher mechanical properties (Currey,1998; Ferguson, 2003; Moffat, 2006; Radhakrishnan, 2004). For example,Ferguson et al. found a positive correlation between indentation modulusand hardness with mineralization in calcified human articular cartilage.Moffat et al. reported that increases in compressive modulus of themineralized fibrocartilage region at the ligament-bone insertioncorresponded to the presence of minerals (Moffat, 2006). Theseobservations collectively suggest that both collagen alignment andmineral content are critical design parameters for interface tissueengineering.

Accordingly, one embodiment of the present invention is an implantabledevice for soft-tissue or soft tissue-to-bone fixation, repair,augmentation, or replacement comprising a biomimetic and biodegradablenanofiber scaffold, which scaffold comprises one or more continuousphases.

Another embodiment of the invention is an implantable biphasicbiomimetic and biodegradable nanofiber device for soft tissue or softtissue-to-bone interface fixation, repair, augmentation, or replacement.This device comprises a first phase comprising nanofibers made from abiodegradable polymer and a second phase coupled to the first phase,which second phase comprises nanofibers made from a biodegradablepolymer and a biocompatible ceramic, wherein the first and second phasesare continuous.

A further embodiment of the invention is an implantable device forfixation, repair, augmentation, or replacement of a rotator cuff or atendon-to-bone interface thereof. This devices comprises a biphasic,biomimetic, and biodegradable nanofiber scaffold having a first phasecomprising nanofibers whose anisotropy mimics that of a tendon andnon-mineralized fibrocartilage, which nanofibers are made from abiodegradable polymer and a second phase coupled to the first phase,which second phase comprises nanofibers whose anisotropy mimics that ofmineralized fibrocartilage and bone, which nanofibers are made from abiodegradable polymer and a biocompatible ceramic, wherein the first andsecond phases are continuous.

Another embodiment of the present invention is a method for fixation of,repairing, augmenting, or replacing a damaged soft tissue or softtissue-to-bone interface in a patient. This method comprises affixing abiomimetic, biodegradable, continuous multi-phasic nanofiber scaffold toa surgically relevant site in order to fixate, repair, augment, orreplace the damaged soft tissue or soft tissue-to-bone interface.

Yet another embodiment of the present invention is a method forfixating, repairing, augmenting, or replacing a damaged rotator cuff ina patient. This method comprises affixing a biomimetic and biodegradablecontinuous multiphase nanofiber scaffold to a surgically relevant sitein order to repair, augment, or replace the damaged rotator cuff.

This application further discloses a fully synthetic scaffold for, e.g.,musculoskeletal tissue engineering.

One embodiment of the present invention is a fully synthetic implantablemulti-phased scaffold. This scaffold comprises, in a single continuousconstruct, a plurality of phases designed to mimic the natural anatomyof a tendon or a ligament.

Another embodiment is a fully synthetic implantable multi-phasedscaffold for ligament repair. This scaffold for ligament repaircomprises (i) a first phase comprising a synthetic graft materialsuitable for implantation into a mammal, the synthetic graft materialdimensioned to have a body with first and second ends; (ii) two secondphases, each second phase comprising microspheres or mesh and having abody and first and second ends, the first end of the respective secondphases disposed at each end of the first phase; and (iii) two thirdphases, each third phase comprising a body and first and second ends,the first end of the respective third phases disposed at the second endof each respective second phase, such that the first phase is separatedfrom the respective third phases by each of the second phases, the thirdphases comprising a material suitable for anchoring the scaffold tobone.

A further embodiment is a fully synthetic implantable multi-phasedscaffold for tendon repair. This scaffold for tendon repair comprises(i) a first phase comprising a synthetic graft material suitable forimplantation into a mammal, the synthetic graft material dimensioned tohave a body with first and second ends; (ii) two second phases, eachsecond phase comprising microspheres or mesh and having a body and firstand second ends, the first end of the respective second phases disposedat each end of the first phase; and (iii) two third phases, each thirdphase comprising a body and first and second ends, the first end of therespective third phases disposed at the second end of each respectivesecond phase, such that the first phase is separated from the respectivethird phases by each of the second phases, the third phases comprising amaterial suitable for anchoring the scaffold to bone.

An additional embodiment is a fully synthetic implantable multi-phasedscaffold for anterior cruciate ligament repair. This scaffold foranterior cruciate ligament repair comprises a first phase comprising asynthetic graft material suitable for implantation into a mammal, thesynthetic graft material dimensioned to have a body with first andsecond ends; (ii) two second phases, each second phase comprisingmicrospheres or mesh and having a body and first and second ends, thefirst end of the respective second phases disposed at each end of thefirst phase; and (iii) two third phases, each third phase comprising abody and first and second ends, the first end of the respective thirdphases disposed at the second end of each respective second phase, suchthat the first phase is separated from the respective third phases byeach of the second phases, the third phases comprising a materialsuitable for anchoring the scaffold to bone.

This application further describes multiphasic apparatuses formusculoskeletal tissue engineering.

A scaffold apparatus, according to one preferred embodiment, ismulti-phasic and can support growth, maintenance and differentiation ofmultiple tissue and cell types. The multi-phasic scaffold apparatus isbiomimetic, biodegradable and/or osteointegrative.

This application also provides a scaffold apparatus for fixingmusculoskeletal soft tissue to bone in a subject, said apparatuscomprising two portions, wherein each portion comprises a scaffold,including first through third phases, wherein (i) the first phasecomprises a material which promotes growth and proliferation offibroblasts, (ii) the second phase adjacent to the first phase comprisesa material which promotes growth and proliferation of chondroblasts, and(iii) the third phase adjacent to the second phase comprises a materialwhich promotes the growth and proliferation of osteoblasts.

This application further provides a scaffold apparatus for fixingmusculoskeletal soft tissue to bone in a subject, said scaffoldapparatus comprising (i) a first phase comprising a material whichpromotes growth and proliferation of fibroblasts, (ii) a second phaseadjacent to the first phase comprising a material which promotes growthand proliferation of chondroblasts, and (iii) a third phase adjacent tothe second phase comprising a material which promotes the growth andproliferation of osteoblasts, wherein a degradable cell barrier isinserted between the adjacent phases.

This application further provides a scaffold apparatus for fixingmusculoskeletal soft tissue to bone in a subject, said scaffoldapparatus comprising (i) a first phase comprising a material whichpromotes growth and proliferation of fibroblasts, (ii) a second phaseadjacent to the first phase comprising a material which promotes growthand proliferation of chondroblasts, and (iii) a third phase adjacent tothe second phase comprising a material which promotes the growth andproliferation of osteoblasts, wherein said first phase of the apparatusis coupled to a soft tissue graft.

This application further provides a scaffold apparatus for fixingmusculoskeletal soft tissue to bone in a subject, said scaffoldapparatus comprising (i) a graft collar and (ii) a polymer-fiber meshcoupled to the graft collar to apply compressive mechanical loading tothe graft collar.

BRIEF DESCRIPTION OF THE FIGURES

FIG. 1 shows the structural properties of the aligned and unalignednanofiber scaffolds. Scanning electron micrographs of A) aligned and B)unaligned nanofiber scaffolds are shown A (2000×, bar=20 μm)

FIG. 2 shows mechanical properties of aligned and unaligned nanofiberscaffolds. A representative stress-strain curve for aligned andunaligned nanofiber scaffolds tested in uniaxial tension is shown.

FIG. 3 shows the effects of nanofiber organization on cell morphology.The top panel shows confocal microscopy of human rotator cufffibroblasts at day 1 and day 14 (20×, bar=100 μm). The bottom panelshows scanning electron micrographs of cells grown on aligned andunaligned nanofiber scaffolds at day 1, and day 14 (1000×, bar=50 μm).The fibroblasts remained viable and grew on both types of substrate overtime, with the rotator cuff cells exhibiting phenotypic elongatedmorphology on the aligned nanofiber scaffolds.

FIG. 4 shows gene expression on aligned and unaligned nanofiberscaffolds over time. Integrin expression differed between the alignedand unaligned groups while types I and III collagen expression weremaintained on both nanofiber scaffold types. The α2 expression wasdetected on the unaligned nanofiber scaffolds at all time pointsevaluated.

FIG. 5 shows quantitative analysis of fibroblast response on aligned andunaligned nanofiber scaffolds at day 1 and day 14 post-plating.Nanofiber organization guided cell attachment and was analyzed based onanalysis of A) Mean Vector Angle (0-horizontally aligned) and B) MeanVector Length (0=random, 1=aligned, *p<0.05). The inserted pictures inFIG. 5A shows confocal images of cell on aligned (top) and unaligned(bottom) nanofibers (60×, bar=50 μm).

FIG. 6 shows cell proliferation and matrix elaboration on aligned andunaligned nanofiber scaffolds. FIG. 6A shows that cells grew on bothtypes of nanofiber scaffolds independent of fiber alignment. FIG. 6Bshows immunohistochemical staining for types I and III collagen (Day 7,20×, bar=100 μm). FIG. 6C demonstrates that matrix production on thenanofiber scaffolds was also guided by nanofiber organization, with analigned collagen I matrix found on the aligned nanofiber scaffolds (Day7, mean angle analysis).

FIG. 7 shows the effects of in vitro culture on nanofiber scaffoldmechanical properties. Mechanical properties decreased due to polymerdegradation, and aligned nanofiber scaffolds were significantly strongerthan unaligned nanofiber scaffolds, FIG. 7A shows ultimate tensilestrength; *:p<0.05; FIG. 7B whose elastic modulus, *:p<0.05 vs.unaligned, and #:p<0.05 vs. unaligned cellular scaffolds; FIG. 7C showsyield strength, *:p<0.05 for aligned vs. as-fabricated aligned nanofiberscaffolds, and #:p<0.05 day 1 vs. day 14.

FIG. 8 shows an arthroscopic image of a torn supraspinatus tendon in theright shoulder, posterior view.

FIG. 9A shows a supraspinatus tendon-to-bone insertion site (100) CT:Tendon, UF: Uncalcified Fibrocartilage, CF: Calcified Fibrocartilage, B:Bone.

FIG. 9B shows another view of the supraspinatus tendon-to-bone insertionsite with the interface regions (calcified and non-calcified)highlighted.

FIG. 10A-E show one embodiment of a clinical application of a biphasicnanofiber scaffold according to the present invention.

FIG. 11A-C show that calcium and phosphorus peaks acquired through EDAXanalysis confirmed incorporation of hydroxyapatite (HA) into the PLGAnanofibers. Fiber roughness was found to increase with increasinghydroxyapatite content (1% (A), 5% (B), and 15% (C) HA) as shown byscanning electron microscopy.

FIG. 12 shows a biphasic nanofiber scaffold according to the presentinvention at different magnifications (A and B) under scanning electronmicroscopy. The nanofiber scaffold was fabricated with Phase Aconsisting of PLGA with 0% HA and Phase B containing 5% HA. The phases(A and B) of FIG. 12 are continuous. When tested in tension (C), theelastic modulus of the biphasic constructs (n=3) was found to besignificantly greater than either Phase A or Phase B alone (n=5).

FIG. 13 shows that increasing HA content in PLGA nanofiber meshes had nosignificant effect on alkaline phosphatase gene expression at days 3 and21 as indicated by semiquantitative analysis of PCR band intensities(n=2).

FIG. 14 shows that (A) matrix and (B) cellular morphology correspondedto fiber alignment on PLGA-HA (5%) nanofiber meshes at day 21 as shownby SEM and confocal fluorescence microscopy, respectively. (C)Extracellular matrix consisted of both type I and type II collagen asshown by immunohistochemistry.

FIG. 15 shows co-culture of fibroblasts (FB) and osteoblasts (OB) on acontinuous biphasic nanofiber scaffold according to the presentinvention. Distinct cellular regions were obtained on the nanofiberscaffold with fibroblasts (green, DiO) and osteoblasts (red, Dil)attached only on Phase A and Phase B, respectively, as indicated byfluorescence confocal microscopy (20×).

FIGS. 16A and B show top-side views of two embodiments of the presentinvention.

FIG. 17 shows top-side views of multiphasic embodiments of the inventionthat are layered—biphasic (A) and triphasic (B). In these embodiments,the phases are layered along a vertical axis (y).

FIG. 18 shows top-side views of multiphasic embodiments of the inventionthat are aligned along a horizontal axis (x)—biphasic (A) and triphasic(B). FIG. 18C shows an expanded view of an embodiment where one of thephases is comprised of multiple layers.

FIG. 19 is a schematic depicting the electrospinning process accordingto the present invention.

FIG. 20 is a series of micrographs showing how nanofiber orientation andalignment change with the drum surface velocity during electrospinning.

FIG. 21A shows a schematic diagram of a scaffold apparatus; 21B shows aschematic diagram of another scaffold apparatus; 21C shows a schematicdiagram of a multi-phased scaffold apparatus.

FIG. 22A shows a flow chart for a method for preparing a scaffold; 22Bshows a flow chart for another method for preparing a multi-phasedscaffold.

FIG. 23A shows a posterior view of an intact bovine anterior cruciateligament (ACL) connecting the femur to the tibia (left); 23B shows anenvironmental scanning electron microscope (ESEM) image of transitionfrom ligament (L) to fibrocartilage (FC) to bone (B) at the ACLinsertion (upper right); 23B shows a histological micrograph of similarACL to bone interface additionally showing mineralized fibrocartilage(MFC) zone (lower right).

FIG. 24A-B show Bovine tibial-femoral joint after ACL and insertion siteextraction (right), ACL and insertion sites after excision.

FIG. 25A shows a scanning electron microscopy (SEM) image of Ca—Pnodules on BG surface (3 days in SBF). Nodules are about 1 μm in sizeinitially, and grew as immersion continued (15,000×).

25B shows an Energy-Dispersive X-ray Analysis (EDXA) spectrum of BGsurfaces immersed in SBF for 3 days. The relative Ca/P ratio isapproximately equal to 1.67.

25C shows a Fourier Transform InfraRed (FTIR) Spectra of bioactive glass(BG) immersed in simulated body fluid (SBF) for up to 7 days, with thepresence of an amorphous calcium phosphate (Ca—P) layer at 1 day, and ofcrystalline layer at 3 days.

FIG. 26A-B show ESEM images of Bovine ACL insertion site (1 and 2),including a cross section of the ACL-femur insertion site, ACL fiber (L)left, fibrocartilage region (FC) middle, and sectioned bone (B) right(FIG. 26A: 250×; FIG. 26B: 500×).

FIG. 27A shows a SEM image of the cross section of the femoral insertionzone, 1000×; 27B shows an EDAX spectrum of the femoral insertion zone.The peak intensities of Ca, P are higher compared to those in ligamentregion.

FIG. 28 shows apparent modulus versus indentation X-position acrosssample.

FIG. 29A-B show X-Ray computed axial tomography (CT) scans of discs madeof poly-lactide-co-glycolide (PLAGA) 50:50 and BG submerged in SBF for 0days (FIG. 29A) and 28 days; FIG. 29B shows the formation of Ca—P overtime.

FIG. 30A shows an SEM image of a PLAGA-BG scaffold; 30B shows an EDAXspectrum of PLAGA-BG immersed in SBF for 14 days.

FIG. 31 shows osteoblast grown on PLAGA-BG, 3 weeks.

FIG. 32 shows a comparison of the expression of type I collagen by humanosteoblast-like cells cultured on the PLAGA-BG composite versus ontissue culture polystyrene (TCPS) controls and on PLAGA alone. Showshigher type I collagen type synthesis on PLAGA-BG.

FIG. 33A shows Alizarin Red (ALZ) stain of ACF fibroblasts, 14 days,20×; 33B shows ALZ stain of interface of ACL fibroblasts andosteoblasts, 14 days, 20×; 33C shows ALZ stain of ACL osteoblasts, 14days, 20×; 33D shows Alkaline Phosphatase (ALP) stain of ACLfibroblasts, 7 days, 32×; 33E shows ALP and4′,6-diamidino-2-phenylindole (DAPI) stain of co-culture of osteoblastsand fibroblasts, 7 days, 32×; and 33F shows APL stain of osteoblasts, 7days, 32×.

FIG. 34A-F show images of multiphase scaffold (FIGS. 34A-34C) andblow-ups of respective sections (FIGS. 34D-34F).

FIG. 35A-35C show multiphasic scaffolds for co-culture of ligamentfibroblasts and osteoblasts. FIGS. 35A and 35B show images of the samescaffold; FIG. 35C is a schematic of scaffold design depicting the threelayers.

FIG. 36A-36D show Micromass co-culture samples of osteoblasts andchondrocytes after 14 days. FIG. 36A show cells stained with hematoxylinand eosin (H&E) stain; FIG. 36B shows cells stained with Alcian blue;FIG. 36C shows the location of Type I collagen; FIG. 36D shows thelocation of Type II collagen (green) and nucleic acids (red).

FIGS. 37A and 37B show reverse transcriptase PCR (RT-PCR) gels for day 7micromass samples. FIG. 37A shows Type X collagen expression. FIG. 37Bshows Type II collagen expression (C: control micromass sample; E:experimental osteoblast-chondrocyte co-culture sample).

FIG. 38A-38B show SEM images of cellular attachment to PLAGA-BG scaffoldafter 30 min; FIG. 38A shows chondrocyte control (2000×); FIG. 38B showsco-culture of osteoblasts and chondrocytes (1500×).

38C-38E show cellular attachment to PLAGA-BG scaffolds; FIG. 38C showschondrocyte control, day 1 (500×); 38D shows co-culture of osteoblastsand chondrocytes, day 1 (500×); FIG. 38E shows co-culture of osteoblastsand chondrocytes, day 7 (750×).

FIG. 39 shows a multi-phased scaffold (for in vitro co-culture ofosteoblast and ligament fibroblast cells).

FIG. 40 shows a schematic diagram depicting a fabrication process of acomposite (PLAGA-BG) of PLAGA and BG, in thin film form and as a 3-D,porous scaffold.

FIG. 41 shows EDXA spectra of the PLAGA-BG composite immersed in SBF for14 days.

FIG. 42A-C show fluorescence microscopy images (day 28, ×10) for PhasesA through C, respectively.

42D-E are images showing extracellular matrix production for Phases Band C, respectively.

FIG. 43A-D show SEM images, in another set of experiments:

A) Phase C, Day 0—×1000;

B) Phase C, Day 28—×1000;

C) Phase A, Day 28—×1000; and

D) Phase B, Day 28—×70.

FIG. 44A-F show fluorescence microscopy images:

A) Phase A, Day 0, ×10;

B) Phase B, Day 0, ×10;

C) Phase C, Day 0, ×10;

D) Phase A, Day 28, ×10;

E) Phase B, Day 28, ×10; and

F) Phase C, Day 28, ×10.

FIG. 45A-C show Trichrome images (Day 0, ×10) of Phase A, Phase B andPhase C, respectively; 45D-E show Picrosirius Red images of Phase B andPhase C, respectively; and 45F shows a von Kossa image of Phase C.

FIG. 46A-F show images of osteoblast and fibroblast in culture, inanother set of experiments:

a) Day 0, 5×;

b) Day 0, 5×;

c) Day 1, 5×;

d) Day 2, 5×;

e) Day 1, 32× (cell contact); and

f) Day 1, 32×.

46G-I show stained images:

a) Live-dead stain of 1 hour sample, 5×;

b) ALP stain of osteoblast and fibroblast, day 2, 20×; and

c) Collagen I staining, day 6, 20×.

FIG. 47 shows a schematic of the experimental design, in another set ofexperiments, for in vitro evaluations of human osteoblasts andfibroblasts co-cultured on multi-phased scaffolds.

FIG. 48A shows a graph which demonstrates cell proliferation in PhasesA, B, and C during 35 days of human hamstring tendon fibroblast andosteoblast co-culture on multiphased scaffolds.

48B-C graphically show mechanical testing data for multiphased scaffoldsseeded with human hamstring tendon fibroblasts and human osteoblastsover 35 days of culture (n=4).

FIG. 49A schematically shows a method for producing multiphasicscaffolds, in another set of experiments. First Ethicon PLAGA mesh iscut into small pieces and inserted into a mold. By applying compressionforce (F) and heating (H) at 150° C. for time (t)=20 hours, the meshsegments are sintered into a mesh scaffold, which is removed from themold. Next, PLAGA microspheres are inserted into the mold, sintered,then removed as a second scaffold. The same process is performed for thePLAGA-BG microspheres. Finally, Phases A and B are joined by solventevaporation, then all three scaffolds are inserted into the mold andsintered together, forming the final multiphasic scaffold.

49B shows a schematic of a fibroblast-osteoblast co-culture experimentaldesign.

FIG. 50A: shows graphically a comparison of microsphere initial mass andfinal mass after undergoing a sintering process.

50B-C: shows graphically scaffold phase thicknesses and diameters, inthe experiments of FIG. 49.

FIG. 51A-B: show graphically mechanical testing data for multiphasedscaffolds seeded with human hamstring tendon fibroblasts and humanosteoblasts over 35 days of culture (n=4). Scaffolds were tested inuniaxial compression. Compressive modulus (A) and yield strength (B)were calculated from the resulting stress-strain curves. Both cellseeded (C) and acellular (AC) scaffolds were examined at days 0, 7, 21,and 35. Scaffold compressive modulus was significantly greater at day 0than for all subsequent time points and groups (p<0.05).

FIG. 52A-D show SEMs of electrospun meshes spun at:

A) 1^(st) gear, 7.4 m/s;

B) 2^(nd) gear, 9.4 m/s;

C) 3^(rd) gear, 15 m/s; and

D) 4^(th) gear, 20 m/s.

52E-F show scanning electron microscopy (SEM) images of a multiphasedscaffold, with 85:15 PLAGA electrospun mesh joined with PLAGA:BGcomposite microspheres.

FIG. 53 schematically shows an exemplary multi-phased scaffold as ahamstring tendon graft collar which can be implemented during ACLreconstruction surgery to assist with hamstring tendon-to-bone healing.

FIG. 54 shows a flow chart for another method for preparing a scaffold.

FIG. 55 shows a comparison of a prior HTG scaffold complex utilizinggraft material from a natural source and a fully synthetic implantablescaffold according to the present invention. As shown, the fullysynthetic implantable scaffold has three phases—Phase A is a syntheticgraft, Phase B disposed on either side of Phase A is made frommicrospheres, and Phase C, disposed on either side of the respectivePhase B material, is made from a composite material.

FIG. 56 shows a top-down view of a fully synthetic implantablemultiphase scaffold according to the present invention. FIG. 56B showsan exploded side view of the scaffold of FIG. 56A.

FIG. 57 shows an elevated side view of another embodiment of a fullysynthetic implantable multi-phase scaffold according to the presentinvention.

FIG. 58 shows an elevated side view of an embodiment of a fullysynthetic implantable multi-phase scaffold according to the presentinvention having a single layer. FIG. 58A shows an elevated side view ofan embodiment of a fully synthetic implantable multi-phase scaffoldaccording to the present invention having multiple layers.

FIG. 59 shows a top-down view of another embodiment of a fully syntheticimplantable multi-phase scaffold according to the present inventionhaving the same phases of FIG. 56, but with mesh dividers between eachphase. FIG. 59B shows an exploded elevated side view of a mesh locatedin between adjacent phases of the scaffold of FIG. 59A. FIG. 59C showsan elevated side view of a mesh of a fully synthetic implantablemulti-phase scaffold according to the present invention having multiplelayers.

FIG. 60 shows an elevated side view of a fully synthetic implantablemulti-phase scaffold according to the present invention in which afourth phase is shown for fixation to a bone.

FIG. 61: I and II: ACL-to-bone insertion (Trichrome, 5×) III: BiomimeticTriphasic scaffold (Ø 7.5×6.5 mm).

FIG. 62 shows potential clinical applications of the triphasic scaffold.

FIG. 63 shows clinical application as a bioactive interference screw.

FIG. 64 shows schematic summary of experimental approach.

FIG. 65: I. Multi-phased scaffold design with nanofiber mesh sinteredbetween phases to localize cell seeding. II. Tracking of fibroblasts(Phase A), chondrocytes (Phase B) and osteoblasts (Phase C) on themulti-phased scaffold (Day 1, 10×). Phase specific cell distribution wasmaintained, which successfully localized fibroblasts (Fb), chondrocytes(CH) and osteoblasts (Ob) on Phase A, B and C, respectively.

FIG. 66: In vivo model. I. Schematic of reconstruction model. II.Reconstruction using flexor tendon graft. III. Bone tunnel formed in thefemur and tibia. IV. Microsphere scaffold inserted into the two bonetunnels.

FIG. 67 shows Experimental design for tracking the three types ofimplanted cell populations in vivo and determining their presence over a4-week implantation period.

FIG. 68 shows Experimental design for interface regeneration on thetri-cultured triphasic scaffold in an intra-articular ACL reconstructionmodel.

FIG. 69 shows a schematic view of a triphasic scaffold with degradablecell barrier inserted between adjacent phases.

FIG. 70 shows a schematic view of a triphasic scaffold coupled to asynthetic graft for a ligament.

FIG. 71 shows a schematic view of a scaffold-mesh apparatus coupled witha soft tissue graft.

FIG. 72 shows in vitro cell culture experimental design for Example 1.6.

FIG. 73 shows the mineral content of the 0% HA, 10% HA and 15% HAscaffold in Example 1.6.

FIG. 74 shows mineral chemistry of incorporated HA in Example 1.6 asdetermined by FTIR.

FIG. 75A shows the surface roughness of the 0% HA, 10% HA and the 15% HAnanofiber-based scaffold in Example 1.6.

FIG. 75B-D show fiber diameter of the 0% HA, 10% HA and the 15% HAnanofiber-based scaffold in Example 1.6.

FIG. 76A-D show the fiber diameter decreasing with increasing mineralcontent in Example 1.6.

FIG. 77A shows tensile modulus decreasing with increasing mineralcontent in Example 1.6.

FIG. 77B shows compressive modulus increasing with increasing mineralcontent in Example 1.6.

FIG. 78A shows that cell morphology is heterogenous with both elongatedand spherical cells observed in Example 1.6.

FIG. 78B shows cell proliferation (Cell number) from Day 1-Day 42 inExample 1.6.

FIG. 79A shows normalized collagen content from Day 7-Day 42 in Example1.6.

FIG. 79B shows chondrocyte collagen deposition with Picrosirius RedStain (20×, bar=100 μm)

FIG. 80 shows nanofiber-guided collagen alignment with Picrosirius RedStain on Day 42 (Picrosirius Red+Polarized Light Microscopy, 20×).

FIG. 81 shows that the nanofiber scaffolds in Example 1.6 supportedcollagen I and II production. (Immunohistochemistry, Day 14, 40×)

FIG. 82A shows the Normalized GAG content from Day 7 to Day 42.

FIG. 82B shows chondrocyte GAG Deposition (Alcian Blue stain, 20×,bar=100 μm)

FIG. 83 shows calcified matrix deposition in Example 1.6.

FIG. 84A shows Gene Expression Relative to GAPDH for Collagen X, MMP-13,Ihh and Runx2 for the 0%, 10% and 15% HA nanofiber scaffolds in Example1.6.

FIG. 84B shows sided by side comparison of collagen (Picrosirius Red),GAG (Alcian Blue), and Mineralization (von Kossa) images for the 15% HAscaffold (14 weeks, 20×).

DETAILED DESCRIPTION

In order to facilitate an understanding of the material which follows,one may refer to Freshney, R. Ian. Culture of Animal Cells—A Manual ofBasic Technique (New York: Wiley-Liss, 2000) for certain frequentlyoccurring methodologies and/or terms which are described therein.

TERMS

The following references provide one of skill with a general definitionof many of the terms used in this invention: Singleton et al.,Dictionary of Microbiology and Molecular Biology (2nd Ed. 1994); TheCambridge Dictionary of Science and Technology (Walker ed., 1988); TheGlossary of Genetics, 5th Ed., R. Rieger et al. (eds.), Springer Verlag(1991); and Hale & Marham, The Harper Collins Dictionary of Biology(1991).

Unless defined otherwise, all technical and scientific terms used hereinhave the meaning commonly understood by a person skilled in the art towhich this invention belongs. However, except as otherwise expresslyprovided herein, each of the following terms, as used in thisapplication, shall have the meaning set forth below.

As used herein, “a single continuous construct” means that each phase is“continuous” with the phase adjacent to it. Thus, in the present fullysynthetic implantable multiphased scaffolds, the interface between onephase and the next is designed, e.g., by sintering and other meansdescribed in more detail below, to mimic the natural anatomicalstructure, e.g., of a tendon or of a ligament, particularly theinsertion sites thereof. In the present invention, the scaffold may havemore than one phase, depending on the anatomical architecture of theligament or tendon to be repaired, fixed, augmented, or replaced. Anexemplary number of phases is from about 1 to about 10, such as forexample, from about 2 to about 4, preferably 3 or 4. As noted above, insuch multiphasic embodiments, each phase of the scaffold is continuousfrom phase-to-phase.

As used herein, “aligned fibers” shall mean groups of fibers which areoriented along the same directional axis. Examples of aligned fibersinclude, but are not limited to, groups of parallel fibers.

As used herein, “bioactive” shall include a quality of a material suchthat the material has an osteointegrative potential, or in other wordsthe ability to bond with bone. Generally, materials that are bioactivedevelop an adherent interface with tissues that resist substantialmechanical forces.

As used herein, a “biocompatible” material is a synthetic or naturalmaterial used to replace part of a living system or to function inintimate contact with living tissue. Biocompatible materials areintended to interface with biological systems to evaluate, treat,augment or replace any tissue, organ or function of the body. Thebiocompatible material has the ability to perform with an appropriatehost response in a specific application and does not have toxic orinjurious effects on biological systems. One example of a biocompatiblematerial can be a biocompatible ceramic.

As used herein, “biodegradable” means that the scaffold, once implantedinto a host, will begin to degrade. The implantable devices of thepresent invention may be used in open surgical procedures, so-called“mini-open” procedures, and arthroscopic procedures as may be requiredand determined by a surgeon. Preferably, the implantable devices of thepresent invention are used in arthroscopic procedures. The nanofiberscaffold of the implantable device is biomimetic and biodegradable. Therate of biodegradation may be engineered into the nanofiber scaffoldbased on the polymers used, the ratio of copolymers used, and otherparameters well known to those of skill in the art. Moreover, in certainembodiments of the present invention, when the nanofiber scaffold is amultiphasic nanofiber scaffold, the rate of biodegradation of each phasemay be separately engineered according to the needs of the particularsurgery to be performed.

As used herein, “biomimetic” shall mean a resemblance of a synthesizedmaterial to a substance that occurs naturally in a human body and whichis not rejected by (e.g., does not cause an adverse reaction in) thehuman body. When used in connection with the “nanofiber scaffold”,biomimetic means that the nanofiber scaffold is biologically inert(i.e., will not cause an immune response/rejection) and is designed toresemble a structure (e.g., soft tissue anatomy) that occurs naturallyin a mammalian, e.g., human, body and that promotes healing whenimplanted into the body.

As used herein, “chondrocyte” shall mean a differentiated cellresponsible for secretion of extracellular matrix of cartilage.Preferably the cells are from a compatible human donor. More preferably,the cells are from the patient (i.e., autologous cells).

As used herein, “fibroblast” shall mean a cell of connective tissue thatsecretes proteins and molecular collagen including fibrillarprocollagen, fibronectin and collagenase, from which an extracellularfibrillar matrix of connective tissue may be formed. Fibroblastssynthesize and maintain the extracellular matrix of many tissues,including but not limited to connective tissue. The fibroblast cell maybe mesodermally derived, and secrete proteins and molecular collagenincluding fibrillar procollagen, fibronectin and collagenase, from whichan extracellular fibrillar matrix of connective tissue may be formed. A“fibroblast-like cell” means a cell that shares certain characteristicswith a fibroblast (such as expression of certain proteins).

As used herein, “fully synthetic” scaffold means that the scaffold iscomposed of man-made material, such as synthetic polymer, or apolymer-ceramic composite, but it does not preclude further treatmentwith material of biological or natural origin, such as seeding withappropriate cell types, e.g., seeding with osteoblasts, osteoblast-likecells, and/or stem cells, or treating with a medicament, e.g.,anti-infectives, antibiotics, bisphosphonate, hormones, analgesics,antiinflammatory agents, growth factors, angiogenic factors,chemotherapeutic agents, anti-rejection agents, and RGD peptides.

As used herein, “functional” shall mean affecting physiological orpsychological functions but not organic structure.

As used herein, “glass transition temperature” means the temperature atwhich, upon cooling, a noncrystalline ceramic or polymer transforms froma supercooled liquid into a rigid glass. The noncrystalline ceramic orpolymer may be of multiple form and composition, and may be formed asmicrospheres. In the context of a sintering process, such as discussedin this application, the polymer chains from adjacent microspherestypically entangle, effectively forming a bond between the microspheresupon cooling. As the polymer is heated above its glass transitiontemperature, long range polymer chain motion begins.

As used herein, “hydrogel” shall mean any colloid in which the particlesare in the external or dispersion phase and water is in the internal ordispersed phase.

As used herein, “graft fixation device” means a device for fixation of agraft, including but not limited to staples, interference screws with orwithout washers, press fit EndoButton® devices and Mitek® anchordevices.

As used herein, “graft” shall mean the device to be implanted duringmedical grafting, which is a surgical procedure to transplant tissuewithout a blood supply, including but not limited to soft tissue graft,synthetic grafts, and the like.

As used herein, “imparted” means treated, including but not limited toapplication of medicament on the surface of the scaffold, integration ofmedicament within the scaffold, or a combination of the two. At leastone of the first, second, or third phases of the synthetic implantablemulti-phased scaffold may also be imparted with a medicament, such asanti-infectives, antibiotics, bisphosphonate, hormones, analgesics,anti-inflammatory agents, growth factors, angiogenic factors,chemotherapeutic agents, anti-rejection agents, or RGD peptides.Preferably, at least one of the first, second, or third phases of thesynthetic implantable multi-phased scaffold is imparted with growthfactors. In this way, delivery of medicaments, particularly growthfactors to, e.g., specific anatomic regions is achievable.

As used herein, “implantable device” according to the present inventionis a surgically appropriate, e.g., biocompatible, apparatus having thedesign and physical properties set forth in more detail below.Preferably, the implantable device is designed and dimensioned tofunction in the surgical repair, augmentation, or replacement of damagedsoft tissue, such as, e.g., a rotator cuff, including fixation oftendon-to-bone. More particularly, the implantable device comprises a“nanofiber scaffold”.

As used herein, “implantable” or “suitable for implantation” meanssurgically appropriate for insertion into the body of a host, e.g.,biocompatible, or having the design and physical properties set forth inmore detail below.

As used herein, “interference screw” means a device indicated for softtissue-bone fixation, specifically, a type of graft fixation devicewhich anchors a flexible transplant like a tendon or a ligament in anopening in a bone. The screw generally has a screw body, a head at oneend of said screw body and a penetrating end at an opposite end of saidscrew body. The device may be used in, for example, anterior cruciateligament surgery. The device may be metallic or bioabsorbable and mayinclude, but is not limited to, titanium cannulated interference screws,Poly-L-Lactide (PLLA) interference screws, etc.

As used herein, “matrix” shall mean a three-dimensional structurefabricated from biomaterials. The biomaterials can bebiologically-derived or synthetic.

As used herein, “mesh” means a network of material. The mesh of thesecond phase may be woven synthetic fibers, non-woven synthetic fibers,and nanofibers suitable for implantation into a mammal, e.g., a human.The woven and non-woven fibers may be made according to well knowntechniques. The nanofiber mesh may be made according to techniques knownin the art and those disclosed in, e.g., co-owned internationalapplication no. PCT/US2008/001889 filed on Feb. 12, 2008 to Lu et al.,which application is incorporated by reference as if recited in fullherein.

As used herein, “microspheres”, i.e., the microspheres of the secondphase, mean microbeads, which are suitable, e.g., for cell attachmentand adhesion. The microspheres of the second phase may be made frompolymers such as aliphatic polyesters, poly(amino acids),copoly(ether-esters), polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(ε-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, or biopolymers, or a blend oftwo or more of the preceding polymers. Preferably, the polymer comprisesat least one of the following materials: poly(lactide-co-glycolide),poly(lactide) or poly(glycolide). More preferably, the polymer ispoly(lactide-co-glycolide) (PLGA).

As used herein, “nanofiber mesh” shall mean a flexible netting ofnanofibers, oriented such that at least some of the nanofibers are notparallel to others of the nanofibers.

As used herein, “nanofiber scaffold” is constructed of “nanofibers.” Asused herein, “nanofiber” shall mean fibers with diameters no more than1000 nanometers. In the present invention, a “nanofiber” is abiodegradable polymer that is electrospun into a fiber as described inmore detail herein below. The nanofibers of the scaffold are oriented insuch a way (i.e., aligned or unaligned) so as to mimic the naturalarchitecture of the soft tissue to be repaired. Moreover, the nanofibersand the subsequently formed nanofiber scaffold are controlled withrespect to their physical properties, such as for example, fiberdiameter, pore diameter, and porosity so that the mechanical propertiesof the nanofibers and nanofiber scaffold are similar to the nativetissue to be repaired, augmented or replaced. Thus, in the case of arotator cuff repair, the nanofiber scaffold is able to regenerate thenative insertion of tendon-to-bone through interface tissue engineeringand promote tendon-to-bone integration and biological fixation. In thepresent invention, such a nanofiber scaffold may be multiphasic, such ase.g., biphasic. One aspect of such multiphasic nanofiber scaffolds isthat each phase is “continuous” with the phase adjacent to it. Thus, inthe present nanofiber scaffolds, the interface between one phase and thenext is designed, e.g., by electrospinning and other means described inmore detail below, to mimic the natural anatomical transition between,e.g., tendon and bone at a tendon-to-bone interface. By designing thenanofiber scaffolds of the present invention so that the phases arecontinuous, improved fixation and function is achieved by minimizingstress concentrations and mediating load transfer between tendon andbone compared to prior systems. In the present invention, the nanofiberscaffold may be engineered to remain in place for as long as thetreating physician deems necessary. Typically, the nanofiber scaffoldwill be engineered to have biodegraded between 6-18 months afterimplantation, such as for example 12 months.

As used herein, “osteoblast” shall mean a bone-forming cell which formsan osseous matrix in which it becomes enclosed as an osteocyte. It maybe derived from mesenchymal osteoprogenitor cells. The term may also beused broadly to encompass osteoblast-like, and related, cells, such asosteocytes and osteoclasts. An “osteoblast-like cell” means a cell thatshares certain characteristics with an osteoblast (such as expression ofcertain proteins unique to bones), but is not an osteoblast.“Osteoblast-like cells” include preosteoblasts and osteoprogenitorcells. Preferably the cells are from a compatible human donor. Morepreferably, the cells are from the patient (i.e., autologous cells).

As used herein, “osteointegrative” means having the ability tochemically bond to bone.

As used herein, “particle reinforcement” means a process for forming acomposite with a higher strength than the original material (forexample, a polymer) by adding particles of a reinforcing material with ahigher strength (for example, a ceramic).

As used herein, “polymer” means a chemical compound or mixture ofcompounds formed by polymerization and including repeating structuralunits. Polymers may be constructed in multiple forms and compositions orcombinations of compositions.

As used herein, “porosity” means the ratio of the volume of intersticesof a material to a volume of a mass of the material.

As used herein, “sintering” shall mean densification of a particulatepolymer compact involving a removal of pores between particles (whichmay be accompanied by equivalent shrinkage) combined with coalescenceand strong bonding between adjacent particles. The particles may includeparticles of varying size and composition, or a combination of sizes andcompositions.

As used herein, “soft tissue graft” shall mean a graft which is notsynthetic, and can include autologous grafts, syngeneic grafts,allogeneic grafts, and xenogeneic graft.

As used herein, “soft tissue” includes, as the context may dictate,tendon and ligament, as well as the bone to which such structures may beattached. Preferably, “soft tissue” refers to tendon- or ligament-boneinsertion sites requiring surgical repair, such as for exampletendon-to-bone fixation.

As used herein, “stem cell” e.g., a mesenchymal stem cell, means anunspecialized cell that has the potential to develop into many differentcell types in the body, such as mesenchymal osteoprogenitor cells,osteoblasts, osteocytes, osteoclasts, chondrocytes, and chondrocyteprogenitor cells. Preferably the cells are from a compatible humandonor. More preferably, the cells are from the patient (i.e., autologouscells).

As used herein, “synthetic graft material”, i.e., first phase, meansman-made material that is intended for insertion into a host body.

The synthetic graft material used in the first phase may be made fromaliphatic polyesters, poly(amino acids), copoly(ether-esters),polyalkylenes oxalates, polyamides, poly(iminocarbonates),polyorthoesters, polyoxaesters, polyamidoesters, poly(ε-caprolactone)s,polyanhydrides, polyarylates, polyphosphazenes, polyhydroxyalkanoates,polysaccharides, degradable polyurethanes, or biopolymers, or a blend oftwo or more of the preceding polymers. Preferably, the synthetic graftmaterial used in the first phase is poly(lactide-co-glycolide),poly(lactide) or poly(glycolide).

As used herein, “synthetic” shall mean that the material is not of ahuman or animal origin.

As used herein, all numerical ranges provided are intended to expresslyinclude at least the endpoints and all numbers that fall between theendpoints of ranges.

The following examples are provided to further illustrate the devicesand methods of the present invention. These examples are illustrativeonly and are not intended to limit the scope of the invention in anyway.

Embodiments

The Biomimetic Nanofiber Scaffold for Soft Tissue and SoftTissue-to-Bone Repair, Augmentation and Replacement

The ideal nanofiber scaffold for rotator cuff tendon repair must be ableto meet the functional demand of the native tendon by matching itsmechanical properties as well as promoting host cell-mediated healing bymimicking the ultrastructure organization of the native tendon. Inaddition to being biomimetic and able to promote cell attachment andgrowth, the nanofiber scaffold must be biodegradable so it can begradually replaced by new tissue without compromising graft mechanicalproperties. To this end, the present invention is directed to ananofiber scaffold for, inter alia, rotator cuff repair, augmentation,or replacement, including fixation of tendon-to-bone. These nanofiberscaffolds are highly advantageous for orthopedic tissue engineering dueto their superior biomimetic potential and physiological relevancebecause they exhibit high aspect ratio, surface area, porosity andclosely mimic the native extracellular matrix (Ma, 2005; Christenson,2007; Pham, 2006; Li, 2007; Murgan, 2007). The biomimetic design of thenanofiber-based scaffold include fiber diameter mimicking collagenfibrils, fiber organization guides cell and ECM alignment, controlledmechanical properties which mimic the native matrix, and high porosityand surfe area-to-volume ratio.

Nanofiber scaffolds have been investigated for bone (34Yoshimoto, 2003;Garreta, 2007; Fujihara, 2005; Badmi, 2006), meniscus (38),intervertebral disk (Nerukar, 2007), cartilage (Li, 2003; Li, 2005),ligament (Lee, 2005; Bashur, 2006) as well as tendon tissue engineering(Sahoo, 2006), and they are likely to be a promising solution for thefunctional augmentation of rotator cuff repairs. Moreover, nanofiberorganization and alignment can be readily modulated during fabrication(Murugan, 2007; Matthews, 2002; Yang, 2005; Pham, 2006). See also, e.g.,U.S. Pat. No. 6,689,166. Thus, nanofiber scaffold systems exhibitsignificant versatility in their ability to tailor structural andmaterial properties to meet the functional demands of, e.g., the rotatorcuff.

Accordingly, one embodiment of the present invention is an implantabledevice for soft-tissue or soft tissue-to-bone fixation, repair,augmentation, or replacement. The device comprises a biomimetic andbiodegradable nanofiber scaffold, which scaffold comprises one or morecontinuous phases. The present invention is well suited for soft-tissuerepairs in mammals, particularly humans. More particularly, theimplantable device comprises a “nanofiber scaffold”.

The implantable devices of the present invention may be used in opensurgical procedures, so-called “mini-open” procedures, and arthroscopicprocedures as may be required and determined by a surgeon. Preferably,the implantable devices of the present invention are used inarthroscopic procedures.

The nanofiber scaffold of the implantable device is biomimetic andbiodegradable. The rate of biodegradation may be engineered into thenanofiber scaffold based on the polymers used, the ratio of copolymersused, and other parameters well known to those of skill in the art.Moreover, in certain embodiments of the present invention, when thenanofiber scaffold is a multiphasic nanofiber scaffold, the rate ofbiodegradation of each phase may be separately engineered according tothe needs of the particular surgery to be performed.

In the present invention, the nanofiber scaffold may be engineered toremain in place for as long as the treating physician deems necessary.Typically, the nanofiber scaffold will be engineered to have biodegradedbetween 6-18 months after implantation, such as for example 12 months.As used herein, all numerical ranges provided are intended to expresslyinclude at least the endpoints and all numbers that fall between theendpoints of ranges.

In one embodiment of the present invention, the nanofiber scaffoldcomprises a plurality of nanofibers that are made from a biodegradablepolymer. In the present invention, the biodegradable polymer may beselected from biodegradable polymer is selected from the groupconsisting of aliphatic polyesters, poly(amino acids), modifiedproteins, polydepsipeptides, copoly(ether-esters), polyurethanes,polyalkylenes oxalates, polyamides, poly(iminocarbonates),polyorthoesters, polyoxaesters, polyamidoesters, poly(ε-caprolactone)s,polyanhydrides, polyarylates, polyphosphazenes, polyhydroxyalkanoates,polysaccharides, modified polysaccharides, polycarbonates,polytyrosinecarbonates, polyorthocarbonates, poly(trimethylenecarbonate), poly(phosphoester)s, polyglycolide, polylactides,polyhydroxybutyrates, polyhydroxyvalerates, polydioxanones, polyalkyleneoxalates, polyalkylene succinates, poly(malic acid), poly(maleicanhydride), polyvinylalcohol, polyesteramides, polycyanoacrylates,polyfumarates, poly(ethylene glycol), polyoxaesters containing aminegroups, poly(lactide-co-glycolides), poly(lactic acid)s, poly(glycolicacid)s, poly(dioxanone)s, poly(alkylene alkylate)s, biopolymers,collagen, silk, chitosan, alginate, and a blend of two or more of thepreceding polymers.

Preferably, the polymer comprises at least one ofpoly(lactide-co-glycolide), poly(lactide), and poly(glycolide). Morepreferably, the polymer is a copolymer, such as for example apoly(D,L-lactide-co-glycolide (PLGA). The advantages of the PLGAnanofiber scaffold include that it is 1) biomimetic and guides tendonregeneration, 2) biodegradable and replaced by host tissue, 3) exhibitphysiologically relevant mechanical properties, and 4) enable biologicalfixation of tendon-to-bone.

As noted above, the ratio of polymers may be varied in the biocompatiblepolymer of the nanofibers in order to achieve certain desired physicalproperties, including e.g., strength, ease of fabrication,degradability, and biocompatibility. Preferably, the ratio of polymersin the biocompatible polymer, e.g., the PLGA copolymer, is between about25:75 to about 95:5. More preferably, the ratio of polymers in thebiocompatible polymer, e.g., the PLGA copolymer, is between about 85:15.Generally, a ratio of about 25:75 in the PLGA copolymer will equate to adegradation time of about six months, a ratio of about 50:50 in the PLGAcopolymer will equate to a degradation time of about twelve months, anda ratio of about 85:15 in the PLGA copolymer will equate to adegradation time of about eighteen months.

As noted above, the anisotropy of the nanofibers in the nanofiberscaffold may be controlled. In the present invention, the anisotropy ofthe nanofibers in the nanofiber scaffold may be varied betweensubstantially aligned to substantially unaligned, depending on theanatomical structure of the soft tissue or soft tissue-to-bone interfaceto be repaired, augmented, fixated, or replaced. For example, thenanofiber alignment and orientation may be designed with reference tothe alignment and orientation of various extracellular matrixcomponents, such as collagen, which as noted above, has an average fiberangle of 83-98° (non-mineralized) and 86-103° (mineralized andfibrocartilage) depending on the region measured (Thomopoulos, 2006). Inone aspect of the invention, it is preferred that the nanofibers arealigned. In another aspect, the nanofibers are unaligned. In a furtheraspect, the nanofiber scaffold may contain regions where the orientationof the nanofibers varies from substantially aligned to substantiallyunaligned. Thus, in one embodiment, the nanofiber scaffold comprisesboth aligned and unaligned nano fibers.

As noted above, the nanofiber alignment and orientation mimics theanatomy of the soft-tissue or soft tissue-to-bone interface to berepaired, augmented, fixated, or replaced. Mimicking the anatomy of softtissue or soft tissue-to-bone interface means having structure anddesign mimicking that of the native soft tissue or soft tissue-to-boneinterface. Preferably, the soft tissue to be repaired, augmented, orreplaced is a ligament or tendon. More preferably, the soft tissue is arotator cuff. For example, the nanofiber alignment and orientationmimics the anatomy of a tendon-to-bone interface, such as, e.g., arotator cuff tendon-to-bone interface. (shown in FIG. 9) Thus, in oneaspect, the nanofiber scaffold is designed to mimic soft tissue andcomprises a preformed interface region. Other soft tissue and softtissue-to-bone interfaces in a mammal, particularly a human, are wellknown to those of skill in the art and are contemplated herein.

In the present invention, the nanofiber scaffold may have one or morephases, depending on the anatomical architecture of the soft tissue orsoft tissue-to-bone interface to be repaired, fixated, augmented, orreplaced. An exemplary number of phases is from about 1 to about 10,such as for example, from about 2 to about 4. In one aspect of theinvention, phases may be adjacent to each other in a single sheet (FIG.18A (biphasic, with phases 70 and 80) and B (triphasic, with phases 90,100, and 110). In this embodiment, each phase is aligned along ahorizontal axis (x). In another aspect, the phases may be layered, oneover another (FIG. 17A (two layers, 20, 30) and B (three layers (40, 50,and 60)). In this embodiment, the phases are aligned/layered along avertical axis (y). In another embodiment, at least two phases arelayered along a vertical axis and at least two phases are aligned alonga horizontal axis. In another aspect, at least one of the phasescomprises more than one layer, e.g., from about 2 to about 20 layers(see, e.g., FIG. 18C insert showing a phase (80) comprised of anexemplary three layers (120, 130, and 140)), and each layer may becomposed of the same or different nanofiber polymer and/or biocompatibleceramic, nanofiber alignment and orientation, and coating. For example,in one biphasic embodiment, a first phase may contain two layers: afirst layer having nanofibers aligned in a parallel arrangement and asecond layer having the nanofibers arranged in an unaligned manner.Similarly, in another biphasic embodiment, a first phase may contain twolayers: a first layer having nanofibers aligned in a perpendiculararrangement and a second layer having the nanofibers arranged in anunaligned manner. Thus, depending on the physical properties desired inthe overall scaffold, one or more layers of nanofibers may be arranged,wherein each layer has the same or different alignment (e.g., parallel,perpendicular, unaligned (or any variation therebetween)). Thus,different properties, particularly viscoelastic responses that mimic thenatural architecture of a native soft tissue or soft tissue-to-boneinterface, may be engineered into the scaffold by, e.g., varying thenumber and alignment of each layer within a particular phase.

Preferably, the nanofiber scaffold is multiphasic, such as for examplebiphasic. As noted above, in such multiphasic embodiments, each phase ofthe scaffold is continuous from phase-to-phase.

Referring now to FIG. 17, in one embodiment of the biphasic design, theimplantable device (15) includes a first phase (20) comprisingnanofibers made from a biodegradable polymer. The implantable deviceincludes a second phase (30), which is coupled to the first phase. Inthe present invention, the phases are coupled to each other usingstandard techniques, such as those disclosed in more detail in theexamples. The second phase comprises nanofibers made from abiodegradable polymer and a biocompatible ceramic. In this embodiment,the first phase is continuous with the second phase.

In this embodiment, the biocompatible ceramic may be incorporated intothe biodegradable polymer by any conventional means. For example, thebiocompatible ceramic may be incorporated into the biodegradable polymerto form a composite nanofiber by solution immersion (Lu, 2003; Lu,2005), liposome delivery, or electrospinning (Wutticharoenmongkol, 2006;Sui, 2007) as described in more details in the Examples. Preferably, thebiocompatible ceramic is incorporated into the nanofibers of the secondphase by electrospinning.

The biocompatible ceramic may be selected from any ceramic material thatis biologically inert (or substantially inert), is incorporatable intothe nanofiber scaffold, and will enhance the nanofiber scaffold'smimicry of mineralized and non-mineralized anatomy in a soft tissue orsoft tissue-to-bone interface to be repaired, fixated, augmented, orreplaced. For example, the biocompatible ceramic may be selected fromsilicon nitride-based ceramics, Pseudowollastonite ceramics (β-CaSiO₃),bredigite (Ca₇MgSi₄O₁₆) ceramics, mono-phase ceramics of monticellite(CaMgSiO(4)), akermanite ceramics (Ca₂MgSi₂O₇), tricalcium silicate(Ca(3)SiO(5)), hydroxyapatite, bio-active glass, calcium phosphate,dense calcium sulfate (DCaS), porous silicated calcium phosphate(Si—CaP), tricalcium phosphate (TCP), calcium pyrophosphate (CPP), andcombinations thereof. Preferably, the biocompatible ceramic ishydroxyapatite or bio-active glass, such as, e.g., 45S5® bioglass(Novabone, Alachua, Fla.).

The biocompatible ceramic may be incorporated into the nanofibers of thescaffold at any convenient concentration based on the method ofincorporation used and the desired physical properties of the scaffold.By way of example, in a composite nanofiber scaffold design according tothe present invention, nanofibers that range from 0 to about 25%biocompatible ceramic may be electrospun. For example, nanofiberscontaining about 1%, about 5%, about 15%, and about 25% hydroxyapatitemay be electrospun.

In another aspect of the invention, a bioactive agent may beincorporated into the nanofiber scaffold of the implantable device. Inthe present invention, the “bioactive agent” may be any pharmaceuticallyacceptable entity that does not deleteriously affect the structure orfunction of the nanofiber scaffold and which may provide an addedbenefit to the patient.

The bioactive agent may be incorporated into a portion or the entiretyof one or more phases (or layers) of the nanofiber scaffolds of thepresent invention. One or more bio-active agents may be distributedthroughout a nanofiber scaffold. In another aspect of the invention, oneor more bioactive agents may be incorporated into a first phase of amultiphasic nanofiber scaffold and one or more other bioactive agents(the same or different from those incorporated into the first phase) maybe incorporated into another phase of the multiphasic nanofiberscaffold. For example, in the case of a biphasic nanofiber scaffold fora rotator cuff repair, a growth factor that promotes growth of tendonfibroblasts may be incorporated into the first phase, which is attachedto the tendon and a growth factor that promotes the growth ofosteoblasts may be incorporated into the second phase, which is attachedto bone.

The bioactive agents may be incorporated into the nanofiber scaffoldusing procedures well known in the art, including, e.g., immersion,impregnation, vacuum suction, spraying, and the like.

Non-limiting representative examples of suitable bioactive agentsaccording to the present invention include an anti-infective, anextracellular matrix component, an antibiotic, bisphosphonate, ahormone, an analgesic, an anti-inflammatory agent, a growth factor, anangiogenic factor, a chemotherapeutic agent, an anti-rejection agent, anRGD peptide, and combinations thereof.

Non-limiting representative examples of suitable growth factorsaccording to the present invention include a member of the TransformingGrowth Factor (TGF) super family, a vascular endothelial growth factor(VEGF), a platelet-derived growth factor (PDGF)₁ an insulin-derivedgrowth factor (IGF), a modulator of a growth factor, and combinationsthereof. In one aspect of this embodiment, a member of the TGF superfamily is selected from TGF-β, bone morphogenetic proteins (BMPs),growth differentiation factors (GDFs), Activin A and Activin B, lnhibinA, lnhibin B, anti-mullerian hormone, Nodal, and combinations thereof.

In another aspect of this embodiment, the TGF-β is selected from TGF-β1,TGF-β2, TGF-β3, and combinations thereof. In a further aspect of thisembodiment, the BMP is selected from the group consisting of BMP1-20 andcombinations thereof. In yet another aspect of this embodiment, the GDFsare selected from GDF1-15 and combinations thereof. In a further aspectof this embodiment, the IGF is selected from IGF1, IGF2, insulin growthfactor binding proteins 1-6 (IGFBP1-6), and combinations thereof. In afurther aspect of this embodiment, a modulator of a growth factor is aSMAD (small mothers against decapentaplegic) selected from SMAD1-9 andcombinations thereof.

The nanofiber scaffold may be treated with other materials to enhance orprovide other additional biological benefits as desired. For example,the nanofiber scaffold may further contain a hydrogel disposed on all ora portion of the scaffold. In a multiphasic nanofiber scaffold, such asfor example, the biphasic nanofiber scaffold, the hydrogel may bedisposed on at least a portion of one or both of the phases (or one ormore layers of a phase).

The hydrogel may be disposed/incorporated into the nanofiber scaffoldusing procedures well known in the art, including immersion,impregnation, vacuum suction, spraying, and the like.

Non-limiting representative examples of suitable hydrogels according tothe present invention are composed of a material selected from agarose,carrageenan, polyethylene oxide, polyethylene glycol, tetraethyleneglycol, triethylene glycol, trimethylolpropane ethoxylate,pentaerythritol ethoxylate, hyaluronic acid, thiosulfonate polymerderivatives, polyvinylpyrrolidone-polyethylene glycol-agar, collagen,dextran, heparin, hydroxyalkyl cellulose, chondroitin sulfate, dermatansulfate, heparan sulfate, keratan sulfate, dextran sulfate, pentosanpolysulfate, chitosan, alginates, pectins, agars, glucomannans,galactomannans, maltodextrin, amylose, polyalditol, alginate-based gelscross-linked with calcium, polymeric chains of methoxypoly(ethyleneglycol)monomethacrylate, chitin, poly(hydroxyalkyl methacrylate),poly(electrolyte complexes), poly(vinylacetate) cross-linked withhydrolysable bonds, water-swellable N-vinyl lactams, carbomer resins,starch graft copolymers, acrylate polymers, polyacrylamides, polyacrylicacid, ester cross-linked polyglucans, and derivatives and combinationsthereof.

In the present invention, hydrogels may contain mammalian cells, suchas, e.g., human cells, in order to promote tissue repair. Non-limitingrepresentative examples of suitable cells that may be incorporated intothe hydrogel and subsequently the nanofiber scaffold includefibroblasts, chondrocytes, osteoblasts, osteoblast-like cells, stemcells, and combinations thereof. Preferably the cells are from acompatible human donor. More preferably, the cells are from the patient(i.e., autologous cells). The cells may be incorporated into a portionor the entirety of one or more phases (or one or more layers of a phase)of the nanofiber scaffolds of the present invention, with or without useof a hydrogel. Moreover, one or more cell types may be distributedthroughout a nanofiber scaffold. In another aspect of the invention, oneor more cell types may be incorporated into a first phase of amultiphasic nanofiber scaffold and one or more other cell types (thesame or different from those incorporated into the first phase) may beincorporated into another phase of the multiphasic nanofiber scaffold.For example, in the case of a biphasic nanofiber scaffold for a rotatorcuff repair, tendon fibroblasts may be incorporated into the firstphase, which is attached to the tendon and osteoblasts may beincorporated into the second phase, which is attached to bone.

In another aspect of the invention, fibroblasts, stem cells,chondrocytes, or combinations thereof are disposed on at least a portionof the first phase of a biphasic nanofiber scaffold. In a furtheraspect, chondrocytes, osteoblasts, osteoblast-like cells, stem cells, orcombinations thereof are disposed on at least a portion of the secondphase of a biphasic nanofiber scaffold. In yet another aspect of thisembodiment, fibroblasts stem cells, and chondrocytes are disposed on atleast a portion of the first phase of a nanofiber scaffold andchondrocytes, osteoblasts, osteoblast-like cells, stem cells, orcombinations thereof are disposed on at least a portion of the secondphase of the nanofiber scaffold.

In one aspect of these embodiments, the stem cells are undifferentiatedprior to disposition on the implantable device. In another aspect, thestem cells are pre-differentiated prior to disposition on theimplantable device. The pre-differentiated stem cells may be selectedfor lineages that are specific for the type of repair to be carried out.For example, stem cells that will differentiate into osteoblasts and/orosteoclasts lineages in the case of a tendon-to-bone interface may beincorporated into the phase of the implantable device that will befixated to bone. Whereas, stem cells that will differentiate into, e.g.,fibroblasts, chondrocytes, and the like may be disposed on the phase ofthe implantable device that will be attached to the tendon.

As disclosed above, the nanofibers and the nanofiber scaffold of theimplantable device are designed to mimic the anatomic architecture ofthe soft tissue or soft tissue-to-bone interface to be repaired,fixated, augmented, or replaced. Thus, the physical and mechanicalproperties of the nanofibers and nanofiber scaffold must approximatethose of the soft tissue or soft tissue-to-bone interface to berepaired, fixated, augmented, or replaced. In the case of rotator cuffrepair, implantable devices of the present invention that are biphasicand biomimetic have been designed and made. As discussed in more detailbelow, these nanofiber scaffolds have physical properties that are thesame as or substantially the same as the in vivo architecture.

For example, a single layer of a scaffold composed of aligned nanofibersaccording to the present invention may have a yield strength of about9.8±1.1 MPa, an elastic modulus of about 341±30 MPa, and an ultimatestress of about 12.0±1.5 MPa. In the case of a scaffold composed ofnon-aligned nanofibers, a single layer of such a nanofiber scaffold mayhave a yield strength of about 2.5±0.4 MPa, an elastic modulus of about107±23 MPa, and an ultimate stress of about 3.7±0.2 MPa.

Moreover, in one implantable device according to the present invention,the nanofiber scaffold is composed of nanofibers with a fiber diameterof between about 568 to about 615 nm, a pore diameter of about 4.2 toabout 4.9 μm, a porosity of about 80.7 to about 81.8%, and apermeability of about 5.7 to about 7.9×10″¹² m⁴/N s. As will beunderstood by one skilled in the art, the nanofiber scaffolds may bedesigned having the physical properties of the soft tissue or softtissue-to-bone interface to be repaired. Thus, the parameters of eachphysical characteristic (e.g., yield strength, elastic modulus, ultimatestress, fiber diameter, pore diameter, and permeability) will bedesigned according to the repair to be carried out. The specific valuesfor these characteristics may be determined from the literature and/orare readily measured using conventional techniques.

As discussed in more detail below, each of these physical properties canbe modified, as desired, to approximate the natural architecture of thesoft tissue to be repaired, augmented, or replaced by making theappropriate selection of polymer and/or polymer ratio, by modifying theelectrospinning process, and by the selection of biocompatible ceramicmaterials and/or hydrogel for incorporation into the nanofiber scaffold.

For example, the orientation and alignment of the nanofibers may bemodified based on whether the nanofibers are spun onto a static surface,which produces fibers of decreased orientation and alignment or whetherthe nanofibers are spun onto a rotating drum. As shown in FIG. 20 anddescribed in more detail in Example 1, increasing drum surface velocityincreases the degree of fiber alignment and orientation.

In another embodiment of the present invention, there is provided animplantable biphasic biomimetic and biodegradable nanofiber device forsoft-tissue or soft tissue-to-bone interface fixation, repair,augmentation, or replacement. This device comprises a first phasecomprising nanofibers made from a biodegradable polymer and a secondphase coupled to the first phase, which second phase comprisesnanofibers made from a biodegradable polymer and a biocompatibleceramic, wherein the first and second phases are continuous. In afurther embodiment of the present invention, there is provided animplantable device for fixation, repair, augmentation, or replacement ofa rotator cuff or a tendon-to-bone interface thereof. This devicecomprises a biphasic, biomimetic, and biodegradable nanofiber scaffoldthat mimics a tendon-to-bone interface. This device has a first phasecomprising nanofibers whose anisotropy mimics that of a tendon andnon-mineralized fibrocartilage, which nanofibers are made from abiodegradable polymer and a second phase coupled to the first phase,which second phase comprises nanofibers whose anisotropy mimics that ofmineralized fibrocartilage and bone, which nanofibers are made from abiodegradable polymer and a biocompatible ceramic, wherein the first andsecond phases are continuous.

In these last two embodiments, the biodegradable polymer is selectedfrom the group consisting of aliphatic polyesters, poly(amino acids),modified proteins, polydepsipeptides, copoly(ether-esters),polyurethanes, polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(ε-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, modified polysaccharides,polycarbonates, polytyrosinecarbonates, polyorthocarbonates,poly(trimethylene carbonate), poly(phosphoester)s, polyglycolide,polylactides, polyhydroxybutyrates, polyhydroxy valerates,polydioxanones, polyalkylene oxalates, polyalkylene succinates,poly(malic acid), poly(maleic anhydride), polyvinylalcohol,polyesteramides, polycyanoacrylates, polyfumarates, poly(ethyleneglycol), polyoxaesters containing amine groups,poly(lactide-co-glycolides), poly(lactic acid)s, poly(glycolic acid)s,poly(dioxanone)s, poly(alkylene alkylate)s, biopolymers, collagen, silk,chitosan, alginate, and a blend of two or more of the precedingpolymers.

In these last two embodiments, the biocompatible ceramic is selectedfrom silicon nitride-based ceramics, Pseudowollastonite ceramics(β-CaSiθ₃), bredigite (Ca₇MgSi₄OiS) ceramics, mono-phase ceramics ofmonticellite (CaMgSiO(4)), akermanite ceramics (Ca₂MgSi₂O₇), tricalciumsilicate (Ca(3)SiO(5)), hydroxyapatite, bio-active glass, calciumphosphate, dense calcium sulfate (DCaS), porous silicated calciumphosphate (Si—CaP), tricalcium phosphate (TCP), calcium pyrophosphate(CPP), and combinations thereof.

In another aspect of these embodiments, at least one of the phasesfurther comprises a bioactive agent selected from an anti-infective, anextracellular matrix component, an antibiotic, bisphosphonate, ahormone, an analgesic, an antiinflammatory agent, a growth factor, anangiogenic factor, a chemotherapeutic agent, an anti-rejection agent, anRGD peptide, and combinations thereof.

In these last two embodiments, non-limiting representative examples ofsuitable growth factors according to the present invention include amember of the Transforming Growth Factor (TGF) super family, a vascularendothelial growth factor (VEGF), a platelet-derived growth factor(PDGF), an insulin-derived growth factor (IGF), a modulator of a growthfactor, and combinations thereof. In one aspect of these embodiments, amember of the TGF super family is selected from TGF-β, bonemorphogenetic proteins (BMPs), growth differentiation factors (GDFs),Activin A and Activin B, lnhibin A, lnhibin B, anti-mullerian hormone,Nodal, and combinations thereof.

In another aspect of these embodiments, the TGF-β is selected fromTGF-β1, TGF-β2, TGF-β3, and combinations thereof. In a further aspect ofthis embodiment, the BMP is selected from the group consisting ofBMP1-20 and combinations thereof. In yet another aspect of theseembodiments, the GDFs are selected from GDF1-15 and combinationsthereof. In a further aspect of these embodiments, the IGF is selectedfrom IGF1, IGF2, insulin growth factor binding proteins 1-6 (IGFBP1-6),and combinations thereof. In a further aspect of these embodiments, amodulator of a growth factor is a SMAD (small mothers againstdecapentaplegic) selected from SMAD1-9 and combinations thereof.

In these last two embodiments, the implantable device may furthercomprise a hydrogel disposed on at least a portion of one or both of thephases (or one or more layers of a phase). In this aspect, the hydrogelis composed of a material selected from agarose, carrageenan,polyethylene oxide, polyethylene glycol, tetraethylene glycol,triethylene glycol, trimethylolpropane ethoxylate, pentaerythritolethoxylate, hyaluronic acid, thiosulfonate polymer derivatives,polyvinylpyrrolidone-polyethylene glycol-agar, collagen, dextran,heparin, hydroxyalkyl cellulose, chondroitin sulfate, dermatan sulfate,heparan sulfate, keratan sulfate, dextran sulfate, pentosan polysulfate,chitosan, alginates, pectins, agars, glucomannans, galactomannans,maltodextrin, amylose, polyalditol, alginate-based gels cross-linkedwith calcium, polymeric chains of methoxypoly(ethyleneglycol)monomethacrylate, chitin, poly(hydroxyalkyl methacrylate),poly(electrolyte complexes), poly(vinylacetate) cross-linked withhydrolysable bonds, water-swellable N-vinyl lactams, carbomer resins,starch graft copolymers, acrylate polymers, polyacrylamides, polyacrylicacid, ester cross-linked polyglucans, and derivatives and combinationsthereof.

In these embodiments, the implantable device may further comprisefibroblasts, chondrocytes, osteoblasts, osteoblast-like cells, stemcells, or combinations thereof. In this aspect, fibroblasts, stem cells,chondrocytes, or combinations thereof are disposed on at least a portionof the first phase. In another aspect, chondrocytes, osteoblasts,osteoblast-like cells, stem cells, or combinations thereof are disposedon at least a portion of the second phase. In a further aspect,fibroblasts, stem cells, and chondrocytes are disposed on at least aportion of the first phase and chondrocytes, osteoblasts,osteoblast-like cells, stem cells, or combinations thereof are disposedon the second phase.

In one aspect of these embodiments, the stem cells are undifferentiatedprior to disposition on the implantable device. In another aspect, thestem cells are pre-differentiated prior to disposition on theimplantable device. The pre-differentiated stem cells may be selectedfor lineages that are specific for the type of repair to be carried out.For example, stem cells that will differentiate into osteoblasts and/orosteoclasts lineages in the case of a tendon-to-bone interface may beincorporated into the phase of the implantable device that will befixated to bone. Whereas, stem cells that will differentiate into, e.g.,fibroblasts, chondrocytes, and the like may be disposed on the phase ofthe implantable device that will be attached to the tendon.

In the present invention, the nanofiber scaffolds of the implantabledevice may be manufactured in manner that is convenient for surgicaldelivery. Preferably, the nanofiber scaffolds are manufactured in amanner that closely mimics the architectural anatomy to be repaired,fixated, augmented, or replaced. Thus, to a certain degree, the softtissue or soft tissue-to-bone interface to be repaired will drive thedimensions of the nanofiber scaffolds.

Referring now to FIG. 16A, and by way of example, for a rotator cuff,the implantable device (1) will be about 0.2 to about 2.0 mm thick (D).The shape of the implantable device is not critical, but should beinformed by surgeon preference. Typically, the implantable device (1)may be about 5.0 cm long (L) and about 5.0 cm wide (W). Alternatively,referring to FIG. 16B, the implantable device (10) may be about 10.0 cmin diameter (D). Preferably, the implantable device will be dimensionedso that it is larger than required for the repair, fixation,augmentation, or replacement, so that the surgeon may adjust thedimensions to fit the particular anatomy of the patient.

Another embodiment of the invention is a method for fixation of,repairing, augmenting, or replacing a damaged soft tissue or softtissue-to-bone interface in a patient. This method comprises affixing abiomimetic, biodegradable continuous multiphasic nanofiber scaffoldaccording to the present invention to a surgically relevant site inorder to repair, fixate, augment, or replace the damaged soft tissue orsoft tissue-to-bone interface.

A further embodiment of the present invention is a method for fixating,repairing, augmenting, or replacing a damaged rotator cuff in a patient.This method comprises affixing a biomimetic and biodegradable continuousmultiphase nanofiber scaffold according to the present invention to asurgically relevant site in order to repair, augment, or replace thedamaged rotator cuff.

In sum, the present invention relates to a biodegradable polymer-basednanofiber scaffold designed for repair, fixation, augmentation, orreplacement of tendon-to-bone insertion site damage, such as, forexample, rotator cuff repair. Data disclosed herein include humanrotator cuff fibroblast response on the degradable nanofiber scaffoldsas well as the effects of nanofiber organization (aligned vs. unalignedfibers) on cell attachment and matrix deposition. The novel nanofiberscaffold of the present invention has been designed to match thestructural and mechanical properties of the rotator cuff tendon.Although not wishing to be bound by a particular theory, it is believedthat fibroblast attachment, morphology and matrix elaboration will beguided by the underlying organization of nanofiber scaffolds.

The Fully Synthetic Implantable Multiphased Scaffold

One embodiment of the present invention is a fully synthetic implantablemulti-phased scaffold. This scaffold comprises, in a single continuousconstruct, a plurality of phases designed to mimic the natural anatomyof a tendon or a ligament. Natural anatomy of the tendon or ligamentmeans the native structure and design of the tendon or ligament, both interms of macroscopic anatomy and microscopic anatomy.

In the present invention, the scaffold may have more than one phase,depending on the anatomical architecture of the ligament or tendon to berepaired, fixed, augmented, or replaced. An exemplary number of phasesis from about 1 to about 10, such as for example, from about 2 to about4, preferably 3 or 4. As noted above, in such multiphasic embodiments,each phase of the scaffold is continuous from phase-to-phase.

In one aspect, the composition of each phase of the fully syntheticimplantable multi-phased scaffold is selected to promote growth andmaintenance of soft tissue and/or soft tissue-to-bone interfaces.

In another aspect, the fully synthetic implantable multi-phased scaffoldis biodegradable.

In the present invention, the polymer may be selected from aliphaticpolyesters, poly(amino acids), modified proteins, polydepsipeptides,copoly(ether-esters), polyurethanes, polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(ε-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, modified polysaccharides,polycarbonates, polytyrosinecarbonates, polyorthocarbonates,poly(trimethylene carbonate)s, poly(phosphoester)s, polyglycolide,polylactides, polyhydroxybutyrates, polyhydroxyvalerates,polydioxanones, polyalkylene oxalates, polyalkylene succinates,poly(malic acid), poly(maleic anhydride), polyvinylalcohol,polyesteramides, polycyanoacrylates, polyfumarates, poly(ethyleneglycol), polyoxaesters containing amine groups,poly(lactide-co-glycolides), poly(lactic acid)s, poly(glycolic acid)s,poly(dioxanone)s, poly(alkylene alkylate)s, biopolymers, or a blend oftwo or more of the preceding polymers.

Referring now to FIGS. 56A and B, in a further aspect of the fullysynthetic implantable multi-phased scaffold, the scaffold comprises (i)a first phase (1, 10) comprising a synthetic graft material suitable forimplantation into a mammal, such as, e.g., a human, the synthetic graftmaterial is dimensioned to have a body (13) with first and second ends(11, 12); (ii) two second phases (2 a, 2 b; 20 a, 20 b), each secondphase comprising microspheres or mesh and having a body (21 a, 21 b) andfirst (22 a, 22 b) and second (23 a, 23 b) ends, the ends (23 a, 22 b)of the respective second phases disposed at each end of the first phase(11, 12); and (iii) two third phases (3 a, 3 b; 30 a, 30 b), each thirdphase having a body (31 a, 31 b) and first and second ends (32 a, 33 a;32 b, 33 b). The first end of the respective third phases (3 a, 3 b; 30a, 30 b) is disposed at an end of each respective second phase (22 a, 23b), such that the first phase (1, 10) is separated from the respectivethird phases (3 a, 3 b; 30 a, 30 b) by each of the second phases (2 a, 2b; 20 a, 20 b), the third phases (3 a, 3 b; 30 a, 30 b) comprising amaterial suitable for anchoring the scaffold to bone. FIG. 55 is agraphic rendering of such a scaffold with phases A, B, and Ccorresponding to phases 1, 2 a, b, and 3 a, b (and 10, 20 a, b, and 30a, b), respectively of FIG. 56.

The shape of the scaffold is not critical, but should be informed bysurgeon preference and the type of procedure. Thus, as shown in FIGS. 55and 57, the scaffold may be cylindrical or, as shown in FIGS. 56 and 58,the scaffold may be rectangular. Such shapes are only exemplary and arenot intended to limit the shape of the scaffold in any way. Preferably,the scaffold will be dimensioned so that it is larger than is requiredfor the specific procedure so that the surgeon may adjust the dimensionsto fit the particular anatomy of the patient.

With reference to FIG. 57, a cylindrical embodiment of the scaffold ofthe present invention (65) is shown. The scaffold comprises (i) a firstphase (40) comprising a synthetic graft material suitable forimplantation into a mammal, such as, e.g., a human, the synthetic graftmaterial is dimensioned to have a body with first and second ends; (ii)two second phases (50 a, 50 b), each second phase comprisingmicrospheres or mesh and having a body and first and second ends, thefirst end of the respective second phases disposed at each end of thefirst phase; and (iii) two third phases (60 a, 60 b), each third phasedisposed at the second end of each respective second phase, such thatthe first phase (40) is separated from the respective third phases (60a, 60 b) by each of the second phases (50 a, 50 b), the third phasescomprising a material suitable for anchoring the scaffold to bone and/orto soft tissue.

In the present invention, the fully synthetic scaffold may have multiplephases and layers, which are designed to mimic the natural architectureof the repair site. Thus, in the present invention, the scaffold mayhave, e.g., from 1 to 10 phases and from 1 to 10 layers. In the presentinvention, when a numerical range is provided, all members of the rangeare intended, including the endpoints. By way of example only, FIG. 58Ashows a fully synthetic multiphasic scaffold (115) having multiplephases (70, 80, 90, 100, and 110). Each phase is designed to mimic aspecific architecture at a repair site. In this embodiment, phase (90)corresponds to a synthetic graft material. Phases (80) and (100) may bethe same or different and may be made of microspheres or mesh. Phases(70) and (110) may be the same or different and may be made of amaterial suitable for anchoring the scaffold to bone and/or to softtissue.

Now referring to FIG. 58B, a fully synthetic scaffold (150) according tothe present invention is exemplified having three layers (120, 130, and140, respectively). Each layer of the scaffold may be the same (i.e.,contain the same phases in the same order) or different (i.e., containthe same phases in a different order) depending on the requirements ofthe surgery. Moreover, the dimensions of each phase may be the same ordifferent and may be the same or different from layer to layer. Asshown, layer 120 is comprised of phases 121 a, b, 122 a, b, 123; layer130 is comprised of phases 131 a, b, 132 a, b, 133; and layer 140 iscomprised of phases 141 a, b, 142 a, b, 143.

Now referring to FIG. 60, another embodiment of the present invention isshown. In this embodiment, a fully synthetic multiphasic scaffold (200)having four phases (190; 200 a and 200 b; 300 a and 300 b; and 160 a and160 b). Each phase is designed to mimic a specific architecture at arepair site. In this embodiment, the first phase (190) corresponds to asynthetic graft material as described herein. The second phase (200 a)and (200 b) may be the same or different and may be made of microspheresor mesh as described herein. The third phase (300 a) and (300 b) may bethe same or different as described herein. The fourth phase (160 a and160 b, respectively) is disposed on either side of the third phase (300a and 300 b, respectively). As shown, the respective first ends of thefourth phase are disposed adjacent to the respective second ends of thethird phase. The fourth phase is made of a material that is dimensionedand suitable for mechanical fixation of each respective fourth phase tobone.

The third phase may be made from a polymer-ceramic composite. Thepolymer of the composite may be selected from polyesters, poly(aminoacids), copoly(ether-esters), polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(ε-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, or biopolymers, or a blend oftwo or more of the preceding polymers. The ceramic of the composite maybe selected from bio-active glass, calcium phosphate, hydroxyapatite, orbeta tricalcium phosphate. Preferably, the polymer ispoly(lactide-co-glycolide) and the ceramic is bio-active glass.

The material used in the fourth phase may be made from aliphaticpolyesters, poly(amino acids), copoly(ether-esters), polyalkylenesoxalates, polyamides, poly(iminocarbonates), polyorthoesters,polyoxaesters, polyamidoesters, poly(ε-caprolactone)s, polyanhydrides,polyarylates, polyphosphazenes, polyhydroxyalkanoates, polysaccharides,degradable polyurethanes, or biopolymers, or a blend of two or more ofthe preceding polymers. Preferably, the synthetic graft material used inthe fourth phase is poly(lactide-co-glycolide), poly(lactide) orpoly(glycolide).

The fourth phase is suitably dimensioned for fixation to bone usingconventionally known surgical techniques for mechanical fixation tobone. As shown in FIG. 60, for example, the fibrous ends of the fourthphase (160 a, 160 b) may be threaded through a hole (170 a, 170 b) inbone (180 a, 180 b), e.g., a femur and tibia in the case of an ACLrepair, and may be secured thereto by e.g., tying, screws, nails, andother similar fixation devices and/or techniques. Ideally, eachrespective fourth phase is longer than required in order to allow thesurgeon to trim it as necessary based on the selected fixation method.

The first phase of the synthetic implantable multi-phased scaffold maybe joined to the respective second phases by sintering or solventevaporation, and the third phases may be joined to the second phases bysintering or solvent evaporation. Additionally, the fourth phases may bejoined to the third phases by sintering or solvent evaporation. Othermethods known in the art for forming a continuous construct are alsowithin the scope of the invention.

In one aspect, at least one of the third phase of the syntheticimplantable multi-phased scaffold is adapted to support growth andmaintenance of bone. Accordingly, the scaffold may be adapted for use inligament repair or replacement, e.g., repair or replacement ofanterior-cruciate ligaments, medial collateral ligaments, lateralcollateral ligaments, posterior cruciate ligaments, cricothyroidligaments, periodontal ligaments, anterior sacroiliac ligaments,posterior sacroiliac ligaments, sacrotuberous ligaments, inferior pubicligaments, superior pubic ligaments, suspensory ligaments of the penis,suspensory ligaments of the breast, volar radiocarpal ligaments, dorsalradiocarpal ligaments, ulnar collateral ligaments, or radial collateralligaments.

The first phase may be seeded with at least one of fibroblasts,fibroblast-like cells, and stem cells. At least one second phase may beseeded with at least one of chondrocytes, fibroblasts, and stem cells.At least one third phase may be seeded with at least one of osteoblasts,osteoblast-like cells, and stem cells.

At least one of the first, second, or third phases of the syntheticimplantable multi-phased scaffold may also be imparted with amedicament, such as anti-infectives, antibiotics, bisphosphonate,hormones, analgesics, anti-inflammatory agents, growth factors,angiogenic factors, chemotherapeutic agents, anti-rejection agents, orRGD peptides.

In another aspect, the fully synthetic implantable multi-phased scaffoldmay be adapted for use in tendon repair or replacement, e.g., repair orreplacement of rotator cuff tendons, elbow tendons, wrist tendons,hamstring tendons, patellar tendons, ankle tendons, foot tendons,supra-spinatus tendon, the Achilles tendon, or the patellar tendon.

Another embodiment is a fully synthetic implantable multi-phasedscaffold for ligament repair. This scaffold for ligament repaircomprises (i) a first phase comprising a synthetic graft materialsuitable for implantation into a mammal, the synthetic graft materialdimensioned to have a body with first and second ends; (ii) two secondphases, each second phase comprising microspheres or mesh and having abody and first and second ends, the first end of the respective secondphases disposed at each end of the first phase; and (iii) two thirdphases, each third phase comprising a body and first and second ends,the first end of the respective third phases disposed at the second endof each respective second phase, such that the first phase is separatedfrom the respective third phases by each of the second phases, the thirdphases comprising a material suitable for anchoring the scaffold tobone.

In one aspect, this scaffold further comprises two fourth phases, eachfourth phase disposed at the second end of each respective third phase,the fourth phases comprising a material dimensioned and suitable formechanical fixation of each respective fourth phase to bone.

In this embodiment, the compositions of each of the phases, as well aspreferred compositions of these phases, are as previously disclosedherein. In one aspect, the first phase may be joined to the respectivesecond phases by sintering or solvent evaporation and the respectivethird phases are joined to the respective second phases by sintering orsolvent evaporation. Additionally, the fourth phases are joined to thethird phases by sintering or solvent evaporation.

A further embodiment is a fully synthetic implantable multi-phasedscaffold for tendon repair. This scaffold for tendon repair comprises(i) a first phase comprising a synthetic graft material suitable forimplantation into a mammal, the synthetic graft material dimensioned tohave a body with first and second ends; (ii) two second phases, eachsecond phase comprising microspheres or mesh and having a body and firstand second ends, the first end of the respective second phases disposedat each end of the first phase; and (iii) two third phases, each thirdphase comprising a body and first and second ends, the first end of therespective third phases disposed at the second end of each respectivesecond phase, such that the first phase is separated from the respectivethird phases by each of the second phases, the third phases comprising amaterial suitable for anchoring the scaffold to bone.

In one aspect, this scaffold further comprises two fourth phases, eachfourth phase disposed at the second end of each respective third phase,the fourth phases comprising a material dimensioned and suitable formechanical fixation of each respective fourth phase to bone.

In this embodiment, the compositions of each of the phases, as well aspreferred compositions of these phases, are as previously disclosedherein. In one aspect of this embodiment, the first phase may be joinedto the respective second phases by sintering or solvent evaporation andthe respective third phases are joined to the respective second phasesby sintering or solvent evaporation. Additionally, the fourth phases arejoined to the third phases by sintering or solvent evaporation.

An additional embodiment is a fully synthetic implantable multi-phasedscaffold for anterior cruciate ligament repair. This scaffold foranterior cruciate ligament repair comprises a first phase comprising asynthetic graft material suitable for implantation into a mammal, thesynthetic graft material dimensioned to have a body with first andsecond ends; (ii) two second phases, each second phase comprisingmicrospheres or mesh and having a body and first and second ends, thefirst end of the respective second phases disposed at each end of thefirst phase; and (iii) two third phases, each third phase comprising abody and first and second ends, the first end of the respective thirdphases disposed at the second end of each respective second phase, suchthat the first phase is separated from the respective third phases byeach of the second phases, the third phases comprising a materialsuitable for anchoring the scaffold to bone.

In one aspect, this scaffold further comprises two fourth phases, eachfourth phase disposed at the second end of each respective third phase,the fourth phases comprising a material dimensioned and suitable formechanical fixation of each respective fourth phase to bone.

In this embodiment, the compositions of each of the phases, as well aspreferred compositions of these phases, are as previously disclosedherein. In one aspect, the first phase may be joined to the respectivesecond phases by sintering or solvent evaporation and the respectivethird phases are joined to the respective second phases by sintering orsolvent evaporation. Additionally, the fourth phases are joined to thethird phases by sintering or solvent evaporation.

The fully synthetic implantable multi-phased scaffold of the presentinvention, as previously disclosed, may further comprise a mesh disposedbetween adjacent phases. The mesh may be made from woven fibers,non-woven fibers, and nanofibers. Additionally, the mesh may becomprised of one or more layers.

Now referring to FIG. 59A, a fully synthetic scaffold according to thepresent invention is exemplified having mesh (4 b, 4 c, 4 a, and 4 drespectively) between adjacent phases of the scaffold (i.e., phases 1and 2 a; 1 and 2 b; 2 a and 3 a; and 2 b and 3 b, respectively). FIG.59B shows an exploded elevated side view of a mesh (4 a) located inbetween phases (2 a and 3 a of FIG. 59A). FIG. 59C shows a mesh (5)having three layers (6, 7, and 8, respectively), but the presentinvention contemplates from about 1 to about 10 layers for a meshdivider. When the mesh is made of electrospun nanofibers, the alignmentand orientation of each mesh layer may be varied as described in, e.g.,Lu et al. international application no PCT/US2008/001889. Each layer ofthe mesh may be the same (i.e., contain the same mesh in the same order)or different (i.e., contain the same mesh in a different order)depending on the requirements of the surgery. Moreover, the dimensionsof each mesh may be the same or different and may be the same ordifferent from layer to layer.

The present invention may optionally include other variations in thephases or scaffolds as set forth below. For example, the scaffolds mayhave a gradient of properties (such as structural properties, porediameter, chemical properties, mechanical properties, etc.), for therepair of musculoskeletal tissue. Such scaffolds are preferablymulti-phased, biodegradable, and osteointegrative.

A scaffold apparatus, according to one preferred embodiment, ismultiphasic, including first, second and third phases, and preferablycan support growth, maintenance and differentiation of multiple tissueand cell types.

The first phase comprises a first material adapted for integration andgrowth (for example, by including one or more osteogenic agents,osteogenic materials, osteoinductive agents, osteoinductive materials,osteoconductive agents, osteoconductive materials, growth factors,chemical factors, etc.) of a first tissue type and is seeded with afirst type of cells (for example, osteoblasts, osteoblast-like cells,stem cells, etc.). The material of the first phase may include, but arenot limited to, microspheres, foams, sponges and any other threedimensional (3-D) scaffold construct consisting of polymer and/orceramic. Polymers may include, but are not restricted to, anybiodegradable polymer such as any of the poly-(α-hydroxy acids), ornatural polymers such as silk, collagen, or chitosan. Ceramics mayinclude but are not limited to bioactive glass, hydroxyapatite, betatricalcium phosphate, or any other calcium phosphate material.

The third phase comprises a second material adapted for integration andgrowth of a second tissue type seeded with a second type of cells (forexample, fibroblasts, chondrocytes, stem cells, etc.). The third phasemay include a composite of materials, including, but not limited to,microspheres, a fiber mesh, degradable polymers, etc.

The second phase is an interfacial zone between the first and thirdphases.

The multiphasic scaffold apparatus preferably has a gradient of calciumphosphate content across the phases, and is preferably biomimetic,biodegradable (that is, each phase is degradable) and/orosteointegrative.

A scaffold apparatus for musculoskeletal tissue engineering, accordingto another embodiment, may include microspheres of selected sizes and/orcomposition. The microspheres may be layered to have a gradient ofmicrosphere sizes and/or compositions. The scaffold may provide afunctional interface between multiple tissue types (for example, softtissue and bone).

FIG. 21A shows schematically a multi-phased scaffold apparatus 10comprising phase A, phase B, and phase C. Phases A-C have a gradient ofproperties. The gradient of properties across phases A-C of the scaffoldmay include mineral content (for example, Ca—P), mechanical properties,chemical properties, structural properties, porosity, geometry, etc. Itshould be apparent to one skilled in the art that although apparatus 10has three phases, the apparatus can be integrated in a scaffold withfour or more phases.

For example, the multi-phased scaffold may contain a gradient of Ca—Pconcentrations. Phase A may be constructed of fiber mesh with alignedfibers and with no Ca—P, phase C may be constructed of polymer-ceramiccomposite with high Ca—P, and phase B may be constructed ofpolymer-ceramic composite with lower Ca—P than phase C.

The scaffold apparatus can promote growth and maintenance of multipletissue types. The scaffold may support growth, maintenance anddifferentiation of multiple tissue and cell types. The multi-phasedscaffold may mimic the inhomogeneous properties of the insertion zonebetween soft tissue and bone, resulting in desired growth, phenotypicexpression, and interactions between relevant cell types.

The phases of the scaffold may be inhomogeneous in properties. Thephases may have zonal differences in mineral content and matrixmorphology designed to mimic the tissue-bone interface and to facilitatethe growth and maintenance of different tissues. The phases may differin morphology. For example, phase A can include a porous fibrous mesh,while phases B and C include microspheres. According to anotherembodiment, the scaffold may include a composite of microspheres and afiber mesh.

The scaffold preferably includes multiple phases. According to oneembodiment, one phase (for example, phase A) supports growth andmaintenance of soft tissue, another phase (for example, phase C)supports growth and maintenance of bone, and a third phase is aninterfacial zone between the first and second phases. The first phasefor supporting growth and maintenance of the soft tissue may be seededwith at least one of fibroblasts, chondrocytes and stem cells. Thesecond phase for supporting growth and maintenance of the bone may beseeded with at least one of osteoblasts, osteoblast-like cells and stemcells. The second phase can contain at least one of osteogenic agents,osteogenic materials, osteoinductive agents, osteoinductive materials,osteoconductive agents, osteoconductive materials, growth factors andchemical factors.

Further, at least one of said first phase and said second phase may beseeded with one or more agents by using a microfluidic system.

The third phase may include some of the microspheres. The third phasecan include a gradient of microsphere sizes and/or a gradient ofmicrosphere compositions. The microspheres in the third phase may bejoined by sintering in at least one stage.

The second phase may include additional microspheres. The second phasecan comprise one of polymeric and composite microspheres including arange of diameters or a gradient of diameter. At least some of themicrospheres of the third phase may be in a first range of sizes, andthe additional microspheres of the second phase may be in a second rangeof sizes lower than the first range of sizes.

The second phase can comprise polymeric hydrogels of one of polyethyleneglycol and hydroxyethyl methacrylate. The hydrogel may comprise one ormore of poly(ethylene glycol), agarose, alginate, 2-hydroxyethylmethacrylate and polyacrylamide. The second phase can comprise collagengels with varied mineral content.

The scaffold may include a composite of microspheres and a fiber mesh.The fiber mesh may be a degradable polymer. For example, the first phasemay include a fiber mesh. The fiber mesh of the first phase and themicrospheres of the third phase may be sintered together. The fiber meshmay be electrospun.

The mesh can include one or more desired agents and/or compound. Forexample, at least one of bioactive agents and peptides may coat thesurface of the mesh. The bioactive agents and peptides can enhancedifferentiation, proliferation and attachment of cells and specific celltypes. Also or alternatively, at least one of bioactive agents andpeptides can directly be incorporated into the mesh.

According to one embodiment, the scaffold may include multiple phasesjoined by a gradient of properties. The multiple phases joined by thegradient of properties may be processed through one or more sinteringstages. The gradient of properties across the multiple phases of thescaffold can include mechanical properties, chemical properties, mineralcontent, structural properties, porosity and/or geometry.

The scaffold apparatus can include plural phases of microspheres. Forexample, a first phase of the microspheres can comprise polymer and asecond phase of the microspheres can comprise one of bioactive glass andcalcium phosphate. Varying concentrations of calcium phosphate can beincorporated into the microspheres. The calcium phosphate can beselected from a group comprising tricalcium phosphate, hydroxyapatite,and a combination thereof. The polymer can be selected from a groupcomprising aliphatic polyesters, poly(amino acids),copoly(ether-esters), polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(c-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, and biopolymers, and a blend oftwo or more of the preceding polymers. The polymer can comprise at leastone of poly(lactide-co-glycolide), poly(lactide) and poly(glycolide).

The microspheres may comprise one or more of CaP, bioactive glass,polymer, etc. The microspheres may be processed through one or moresintering stages.

The microspheres may comprise one or more desired agents or compounds.For example, at least one of bioactive agents and peptides may coat thesurface of at least some of the microspheres. The bioactive agents andpeptides can enhance at least one of differentiation, proliferation andattachment of cells and specific cell types. Also or alternatively, atleast one of bioactive agents and peptides can directly be incorporatedinto at least some of the microspheres. The microspheres canadditionally include one or more agents selected from a group comprisingantiinfectives, hormones, analgesics, anti-inflammatory agents, growthfactors, chemotherapeutic agents, anti-rejection agents and RGDpeptides.

The apparatus is preferably biomimetic, biodegradable and/orosteointegrative.

According to one exemplary embodiment, the apparatus may be integratedin a graft fixation device. The graft fixation device may be used, forexample, for graft fixation at the bone tunnels during anterior cruciateligament reconstruction. According to another embodiment, the apparatusmay be integrated in an interference screw.

In addition, the scaffold apparatus, according to another exemplaryembodiment, may be integrated in a graft collar. The graft collar hasmany applications. For example, the graft collar may be adapted forhamstring tendon-to-bone healing. As another example, the graft collarcan be adapted for peridontal ligament repair. Further, the graft collarmay be adapted for spinal repair.

A scaffold apparatus for soft tissue-to-bone interface tissueengineering, according to another exemplary embodiment, is shownschematically in FIG. 21B. Apparatus 15 includes a first region H, asecond region I which is joined to region H, a third region J which isjoined to region I, and a fourth region K which is joined to region J.

Region H comprises composite microspheres of a first size andcomposition optimized to promote growth, proliferation, anddifferentiation of a first cell type for integration and growth of afirst tissue type. The composite microspheres of region H can include arange of sizes.

Region I comprises at least one of microspheres and a fibrous meshhaving a second size and a second composition. The microspheres and/orfibrous mesh of region I can include a range or gradient of sizes,and/or a gradient of compositions.

Region J comprises at least one of a microsphere and a fibrous meshhaving a third size and a third composition. Regions I and J areoptimized to promote growth, proliferation and differentiation of asecond cell type for integration and formation of a second tissue type.The microspheres and/or fibrous mesh of region J can include a range orgradient of sizes, and/or a gradient of compositions.

Region K comprises at least one of a microsphere and a fibrous meshhaving a composition adapted to promote growth, proliferation, anddifferentiation of a third cell type for integration and growth of athird tissue

The regions H-K can be joined together through one of a solid statesintering process and a solvent aggregation process, in which selectedgrowth factors or bioactive agents are incorporated into each region topromote formation, growth and integration of said first, second andthird types of tissues. The scaffold apparatus may be integrated in agraft collar.

A multi-phased scaffold apparatus for providing a functional interfacebetween bone and soft tissue, according to an embodiment schematicallyshown in FIG. 21C, includes microspheres as one phase, and a mesh asanother phase. The microspheres and the mesh may be sintered together.

FIG. 21C shows schematically a multi-phased scaffold apparatus 20comprising phase X and phase Y. Microspheres may be one phase of thescaffold and a mesh may be another phase of the scaffold. Themicrospheres and the mesh may be sintered together. The apparatus 20 maybe integrated in a scaffold which includes multiple phases (for example,three or more).

The microsphere and mesh structure of the scaffold may be geometricallyheterogeneous, including a fiber mesh for culturing fibroblasts and anopen-pore structure for osteoblasts. At least one zone of hydrogels oropen-pore structure for chondrocytes may also be included. Themicrosphere and mesh components may be incorporated into the scaffold toallow for the co-culturing of multiple cell types to mimic the multitudeof cell types found at native tissue interfaces. The mesh can beelectrospun.

The scaffold may be modified to achieve specific cell cultureparameters, for example, by including microspheres of varying diametersto vary the porosity of the scaffold in different regions. Furthermore,the scaffold may be fabricated in a variety of geometries. For example,the scaffold apparatus can be integrated in a graft collar.

This application also describes methods for preparing a scaffold formusculoskeletal tissue engineering. A method, according to oneembodiment (FIG. 22A), includes (a) processing a plurality ofmicrospheres (step S31), including incorporating calcium phosphate intothe microspheres, (b) laying the processed microspheres in a mold (stepS33), the microspheres in the mold presenting a gradient of microspheresizes and/or compositions, and (c) sintering together the microspheresin the mold above a glass transition temperature (step S35).

Additional steps may optionally be added to the method to impartadditional scaffold features or characteristics. For example, the methodmay further include sintering a fiber mesh onto the microsphereconstruct to provide a functional interface between multiple tissuetypes. Further, the method may further comprise electrospinning saidfiber mesh prior to attaching the electrospun fiber mesh onto themicrosphere construct.

Varying concentrations of calcium phosphate may be incorporated into themicrospheres. The calcium phosphate incorporated into the microspheresmay include hydroxyapatite, tricalcium phosphate, etc.

The particulate phase of the microspheres may include bioactive glass.Varying porosity or concentrations of bioactive glass may beincorporated into the microspheres.

The method may further include applying a particle reinforcement processto the microspheres. The method may further include incorporatingparticulates in the microspheres prior to the sintering step tostrengthen the microspheres.

A method for preparing a multi-phase scaffold for musculoskeletal tissueengineering, according to an exemplary embodiment (FIG. 22B), includes(a) processing a plurality of microspheres (step S41), includingincorporating calcium phosphate into the microspheres, (b) laying theprocessed microspheres in a mold (step S43), wherein the microspheres inthe mold presenting a gradient of microsphere sizes for a first phaseand a second phase of the multi-phase scaffold, with microspheres of thefirst phase being in a first range of sizes, and with microspheres ofthe second phase being in a second range of sizes larger than the firstrange of sizes, (c) sintering together the microspheres in the moldabove a glass transition temperature (step S45), and (d) sintering afiber mesh, as a third phase of the multi-phase scaffold, onto themicrosphere construct prepared in (c) (step S47).

Additional steps may optionally be included. For example, the method mayfurther include seeding the third phase with at least one of fibroblasts(for example, human hamstring tendon fibroblasts), chondrocytes and stemcells. The seeding of the third phase supports growth and maintenance ofsoft tissue. Also, the method can include seeding the first phase withat least one of osteoblasts and stem cells. The seeding of the firstphase supports growth and maintenance of bone. The method may furtherinclude seeding the second phase with at least one of chondrocytes andstem cells. Seeding of the second phase can support growth andmaintenance of fibrocartilage.

The first phase may support growth and maintenance of bone. The thirdphase may support growth and maintenance of soft tissue. The secondphase may serve at least as an interfacial zone between the first phaseand the third phase.

For example, the method may further comprise seeding the first phasewith first cells, for supporting growth and maintenance of the bone,seeding the third phase with second cells for supporting growth andmaintenance of the soft tissue, and allowing at least some of said firstcells and said second cells to migrate to the second phase.

In addition, the method may further comprise seeding at least one ofsaid first, second and third phases with one or more agents by using amicrofluidic system.

Further, the method may further comprise electrospinning said fiber meshprior to attaching the fiber mesh onto the microsphere construct.

This application also provides methods for producing polymer/ceramiccomposite microspheres. The composite microspheres can be formed byapplying an emulsion and solvent evaporation process. The compositemicrospheres can comprise a degradable polymer and one of bioactiveglass and calcium phosphate ceramics. The degradable polymer can bedissolved in a solvent. The bioactive glass and/or calcium phosphateceramics can be mixed into the polymer solution. A suspension of thebioactive glass and/or calcium phosphate ceramics in the polymersolution can be poured into a stirring surfactant solution.

Calcium phosphate and/or bioactive glass particles may be encapsulatedin the microspheres during emulsion.

A method, according to another exemplary embodiment (FIG. 54), forpreparing a multi-phase scaffold for musculoskeletal tissue engineering,can comprise the steps of (a) forming a mesh scaffold by sinteringtogether a plurality of mesh segments as a first phase of themulti-phase scaffold (step S351), (b) forming a second scaffold bysintering together a plurality of poly-lactide-co-glycolide microspheresas a second phase of the multi-phase scaffold (step S352), (c) forming athird scaffold by sintering together a plurality of microspheres formedof a composite of poly-lactide-co-glycolide and bioactive glass as athird phase of the multi-phase scaffold (step S353), and (d) sinteringtogether said mesh scaffold, said second scaffold and said thirdscaffold (step S354). Steps S351 through S353 may be performed in anyorder.

The Triphasic Scaffold for Use as Graft Collar or Interference Screw

In an exemplary embodiment (FIG. 62), a scaffold apparatus for fixingmusculoskeletal soft tissue to bone in a subject, comprises twoportions, each portion including first through third phases, wherein (i)the first phase of the scaffold comprises a material which promotesgrowth and proliferation of fibroblasts, (ii) the second phase adjacentto the first phase comprises a material which promotes growth andproliferation of chondroblasts, and (iii) the third phase adjacent tothe second phase comprises a material which promotes the growth andproliferation of osteoblasts.

In the exemplary embodiment of FIG. 62, the two portions (for example,portions 31 and 32) encase respective portions of a soft tissue graft onall sides (for example, halves 31 a and 31 b) of the scaffold apparatus.

In another embodiment (FIG. 63), two portions combine to encase aportion (35) of a soft tissue graft on all sides.

In another embodiment (for example, FIG. 62), the first phase is exposedto the joint cavity. In another embodiment (for example FIG. 62) thesecond phase contacts articular cartilage. In another embodiment (forexample, FIG. 62), the third phase is encased in bone. In anotherembodiment, the interference screw is biomimetic. In another embodiment,the interference screw is biodegradable. In another embodiment, theinterference screw is osteointegrative. In another embodiment, adegradable cell barrier is inserted between the adjacent phases. Inanother embodiment, the degradable cell barrier comprises a nanofibermesh. In another embodiment, the nanofiber mesh comprisespolylactide-co-glycolide (PLGA). In another embodiment, the nanofibermesh is electrospun.

The application further provides an interference apparatus comprising ascaffold apparatus for fixing musculoskeletal soft tissue to bone in asubject, said apparatus comprising two portions, wherein each portioncomprises a scaffold, including first through third phases, wherein (i)the first phase comprises a material which promotes growth andproliferation of fibroblasts, (ii) the second phase adjacent to thefirst phase comprises a material which promotes growth and proliferationof chondroblasts, and (iii) the third phase adjacent to the second phasecomprises a material which promotes the growth and proliferation ofosteoblasts. In one embodiment, the interference apparatus is aninterference screw.

In another exemplary embodiment (FIG. 69), a scaffold apparatus forfixing musculoskeletal soft tissue to bone in a subject, comprises (i) afirst phase (91) comprising a material which promotes growth andproliferation of fibroblasts, (ii) a second phase (92) adjacent to thefirst phase comprising a material which promotes growth andproliferation of chondroblasts, and (iii) a third phase (93) adjacent tothe second phase (92) comprising a material which promotes the growthand proliferation of osteoblasts, wherein a degradable cell barrier (94a, 94 b) is inserted between adjacent phases of the scaffold apparatus.

In one embodiment, the first phase is for supporting growth andmaintenance of soft tissue, the second phase is for supporting thegrowth and maintenance of fibrocartilage, and the third phase is forsupporting the growth and maintenance of bone tissue. In anotherembodiment, the first phase is seeded with at least one of fibroblastsand stem cells. In another embodiment, the stem cells are mesenchymalstem cells. In another embodiment, the first phase includes fiber mesh.In another embodiment, the fiber mesh is electro spun.

In another embodiment, the second phase is seeded with at least one ofchondrocytes and stem cells. In another embodiment, the stem cells aremesenchymal stem cells.

In another embodiment, the third phase is seeded with at least one ofosteoblasts, osteoblast-like cells, and stem cells. In anotherembodiment, the stem cells are mesenchymal stem cells. In anotherembodiment, said third phase contains at least one of osteogenic agents,osteogenic materials, osteoinductive agents, osteoinductive materials,osteoconductive agents, osteoconductive materials, growth factors andchemical factors.

In one embodiment, said scaffold apparatus is integrated in a graftcollar. In another embodiment, said graft collar is adapted forhamstring tendon-to-bone healing. In another embodiment, said firstphase is seeded with human hamstring tendon fibroblasts. In anotherembodiment, said graft collar is adapted for peridontal ligament repair.In another embodiment, said graft collar is adapted for spinal repair.In another embodiment, at least one of said first phase and said thirdphase is seeded with one or more agents by using a microfluidic system.

In one embodiment, the scaffold has multiple phases joined by a gradientof properties. In another embodiment, the multiple phases of thescaffold are processed through one or more sintering stages. In anotherembodiment, the gradient of properties across the multiple phases of thescaffold includes mechanical properties. In another embodiment, thegradient of properties across the multiple phases of the scaffoldincludes chemical properties. In another embodiment, the gradient ofproperties across the multiple phases of the scaffold includes mineralcontent. In another embodiment, the gradient of properties across themultiple phases of the scaffold includes structural properties. Inanother embodiment, the gradient of properties across the multiplephases of the scaffold includes porosity. In another embodiment, thegradient of properties across the multiple phases of the scaffoldincludes geometry.

In one embodiment, the first phase comprises polymer and the third phasecomprises one of bioactive glass and calcium phosphate. In anotherembodiment, the calcium phosphate is selected from a group comprisingtricalcium phosphate, hydroxyapatite, and a combination thereof. Inanother embodiment, the polymer is selected from a group comprisingaliphatic polyesters, poly(amino acids), copoly(ether-esters),polyalkylenes oxalates, polyamides, poly(iminocarbonates),polyorthoesters, polyoxaesters, polyamidoesters, poly(ε-caprolactone)s,polyanhydrides, polyarylates, polyphosphazenes, polyhydroxyalkanoates,polysaccharides, and biopolymers, and a blend of two or more of thepreceding polymers. In another embodiment, the polymer comprises atleast one of poly(lactide-co-glycolide), poly(lactide) andpoly(glycolide).

In one embodiment, the apparatus is biomimetic. In another embodiment,the apparatus is biodegradable. In another embodiment, the apparatus isosteointegrative.

In one embodiment, the apparatus additionally include one or more agentsselected from a group comprising antiinfectives, hormones, analgesics,anti-inflammatory agents, growth factors, chemotherapeutic agents,anti-rejection agents and RGD peptides.

In one embodiment, the degradable cell barrier is a nanofiber mesh. Inanother embodiment, the nanofiber mesh comprisespolylactide-co-glycolide (PLGA). In another embodiment, the nanofibermesh is electrospun.

In another exemplary embodiment, a scaffold apparatus for fixingmusculoskeletal soft tissue to bone in a subject, comprises (i) a firstphase (101) comprising a material which promotes growth andproliferation of fibroblasts, (ii) a second phase (102) adjacent to thefirst phase comprising a material which promotes growth andproliferation of chondroblasts, and (iii) a third phase (103) adjacentto the second phase comprising a material which promotes the growth andproliferation of osteoblasts, wherein said first phase (101) of theapparatus is coupled to a soft tissue graft (104). In one embodiment,the soft tissue graft is a graft of a ligament and the ligament is ananterior cruciate ligament.

A scaffold apparatus, according to one preferred embodiment, ismultiphasic, including phases A, B, and C, and preferably can supportgrowth, maintenance and differentiation of multiple tissue and celltypes.

The Phase A comprises a first material adapted for integration andgrowth of a second tissue type seeded with a second type of cells (forexample, fibroblasts, chondrocytes, stem cells, etc.). Phase A mayinclude a composite of materials, including, but not limited to,microspheres, a fiber mesh, degradable polymers, etc.

Phase C comprises a second material adapted for integration and growth(for example, by including one or more osteogenic agents, osteogenicmaterials, osteoinductive agents, osteoinductive materials,osteoconductive agents, osteoconductive materials, growth factors,chemical factors, etc.) of a first tissue type and is seeded with afirst type of cells (for example, osteoblasts, osteoblast-like cells,stem cells, etc.). The material of the first phase may include, but isnot limited to, microspheres, foams, sponges and any other threedimensional (3-D) scaffold construct consisting of polymer and/orceramic. Polymers may include, but is not restricted to, anybiodegradable polymer such as any of the poly-(α-hydroxy acids), ornatural polymers such as silk, collagen, or chitosan. Ceramics mayinclude but are not limited to bioactive glass, hydroxyapatite, betatricalcium phosphate, or any other calcium phosphate material.

Phase B is an interfacial zone between the first and third phases. Inone embodiment, Phase B is seeded with chondrocytes, such that afibrocartilage interface can be formed and maintained with interactionsbetween these three cell types.

The multiphasic scaffold apparatus preferably is preferably biomimetic,biodegradable (that is, each phase is degradable) and/orosteointegrative.

The scaffold may provide a functional interface between multiple tissuetypes (for example, soft tissue and bone).

FIG. 62 shows an example of a multi-phased scaffold apparatus in theform of a graft collar comprising phase A, phase B, and phase C. Itshould be apparent to one skilled in the art that although the apparatusshown in FIG. 62 has three phases, the apparatus can be integrated in ascaffold with four or more phases.

The scaffold apparatus can promote growth and maintenance of multipletissue types. The scaffold may support growth, maintenance anddifferentiation of multiple tissue and cell types. The multi-phasedscaffold may mimic the inhomogeneous properties of the insertion zonebetween soft tissue and bone, resulting in desired growth, phenotypicexpression, and interactions between relevant cell types.

The phases of the scaffold may be inhomogeneous in properties. Thephases may have zonal differences in mineral content and matrixmorphology designed to mimic the tissue-bone interface and to facilitatethe growth and maintenance of different tissues. The phases may differin morphology. For example, phase A can include a porous fibrous mesh,while phases B and C include microspheres. According to anotherembodiment, the scaffold may include a composite of microspheres and afiber mesh.

The scaffold preferably includes multiple phases. According to oneembodiment, one phase (for example, phase A) supports growth andmaintenance of soft tissue, another phase (for example, Phase B) is aninterfacial zone between the first and second phases and another phase(for example, phase C) supports growth and maintenance of bone. Phase Afor supporting growth and maintenance of the soft tissue may be seededwith at least one of fibroblasts, chondrocytes and stem cells. Phase Cfor supporting growth and maintenance of the bone may be seeded with atleast one of osteoblasts, osteoblast-like cells and stem cells. Phase Ccan contain at least one of osteogenic agents, osteogenic materials,osteoinductive agents, osteoinductive materials, osteoconductive agents,osteoconductive materials, growth factors and chemical factors.

Further, at least one of said Phase A and Phase C may be seeded with oneor more agents by using a microfluidic system.

The scaffold may include a composite of microspheres and a fiber mesh.The fiber mesh may be a degradable polymer. For example, the first phasemay include a fiber mesh. The fiber mesh of the first phase and themicrospheres of the third phase may be sintered together. The fiber meshmay be electrospun.

The mesh can include one or more desired agents and/or compound. Forexample, at least one of bioactive agents and peptides may coat thesurface of the mesh. The bioactive agents and peptides can enhancedifferentiation, proliferation and attachment of cells and specific celltypes. Also or alternatively, at least one of bioactive agents andpeptides can directly be incorporated into the mesh.

According to one embodiment, the scaffold may include multiple phasesjoined by a gradient of properties. The multiple phases joined by thegradient of properties may be processed through one or more sinteringstages. The gradient of properties across the multiple phases of thescaffold can include mechanical properties, chemical properties, mineralcontent, structural properties, porosity and/or geometry.

The scaffold apparatus can include plural phases of microspheres. Forexample, Phase A of the microspheres can comprise polymer and Phase C ofthe microspheres can comprise one of bioactive glass and calciumphosphate. Varying concentrations of calcium phosphate can beincorporated into the microspheres. The calcium phosphate can beselected from a group comprising tricalcium phosphate, hydroxyapatite,and a combination thereof. The polymer can be selected from a groupcomprising aliphatic polyesters, poly(amino acids),copoly(ether-esters), polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(ε-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, and biopolymers, and a blend oftwo or more of the preceding polymers. The polymer can comprise at leastone of poly(lactide-co-glycolide), poly(lactide) and poly(glycolide).

The microspheres may comprise one or more of CaP, bioactive glass,polymer, etc. The microspheres may be processed through one or moresintering stages.

The microspheres may comprise one or more desired agents or compounds.For example, at least one of bioactive agents and peptides may coat thesurface of at least some of the microspheres. The bioactive agents andpeptides can enhance at least one of differentiation, proliferation andattachment of cells and specific cell types. Also or alternatively, atleast one of bioactive agents and peptides can directly be incorporatedinto at least some of the microspheres. The microspheres canadditionally include one or more agents selected from a group comprisinganti-infectives, hormones, analgesics, anti-inflammatory agents, growthfactors, chemotherapeutic agents, anti-rejection agents and RGDpeptides.

The apparatus is preferably biomimetic, biodegradable and/orosteointegrative.

According to one exemplary embodiment, the apparatus may be integratedin a graft fixation device. The graft fixation device may be used, forexample, for graft fixation at the bone tunnels during anterior cruciateligament reconstruction.

According to another embodiment, the apparatus may be integrated in aninterference screw.

In addition, the scaffold apparatus, according to another exemplaryembodiment, may be integrated in a graft collar. The graft collar hasmany applications. For example, the graft collar may be adapted forhamstring tendon-to-bone healing. As another example, the graft collarcan be adapted for peridontal ligament repair. Further, the graft collarmay be adapted for spinal repair.

This application further provides a scaffold apparatus for fixingmusculoskeletal soft tissue to bone in a subject, said scaffoldapparatus comprising (i) a graft collar and (ii) a polymer-fiber meshcoupled to the graft collar to apply compressive mechanical loading tothe graft collar.

In one embodiment, the polymer-fiber mesh wraps around the graft collar.In another embodiment, the surface of the graft collar is wrapped in itsentirety. (FIG. 71)

In one embodiment, the graft collar is biphasic. In another embodiment,the polymer-fiber mesh comprises nanofibers. In yet another embodiment,the nanofiber mesh comprises polylactide-co-glycolide (PLGA). In yetanother embodiment, the nanofiber mesh is electro spun.

According to one embodiment, the scaffold apparatus is coupled to a softtissue graft. (FIG. 71) According to another embodiment, the soft tissuegraft is graft for a ligament. In another embodiment, the ligament is ananterior cruciate ligament.

In another exemplary embodiment (FIG. 71), an interference apparatuscomprising a scaffold apparatus is provided for fixing musculoskeletalsoft tissue to bone in a subject, said scaffold apparatus comprising (i)a graft collar 101 and (ii) a polymer-fiber mesh (105) coupled to thegraft collar to apply compressive mechanical loading to the graftcollar. In one embodiment, the interference apparatus is an interferencescrew.

In one embodiment, the polymer-fiber mesh wraps around the graft collar.In another embodiment, an outer surface of the graft collar is wrappedin its entirety by the polymer-fiber mesh.

In one embodiment, the graft collar is biphasic. In another embodiment,the biphasic graft collar includes a first phase comprising a materialwhich promotes growth and proliferation of fibroblasts, and a secondphase adjacent to the first phase comprising a material which promotesthe growth and proliferation of osteoblasts.

In another embodiment, the polymer-fiber mesh comprises nanofibers. Inyet another embodiment, the nanofiber mesh comprisespolylactide-co-glycolide (PLGA). In yet another embodiment, thenanofiber mesh is electrospun.

In one embodiment, the scaffold apparatus is coupled to a soft tissuegraft. In another embodiment, the soft tissue graft is a graft for aligament of the subject. In yet another embodiment, the ligament is ananterior cruciate ligament of the subject.

This application further provides a graft-fixation apparatus comprisinga scaffold apparatus for fixing musculoskeletal soft tissue to bone in asubject comprising (i) a graft collar and (ii) a polymer-fiber meshcoupled to the graft collar to apply compressive mechanical loading tothe graft collar.

In an exemplary embodiment, the graft fixation apparatus is aninterference screw.

This application further provides a scaffold apparatus for fixingmusculoskeletal soft tissue to bone, said scaffold apparatus beingconfigured to apply mechanical loading to a soft tissue graft to promoteregeneration of a fibrocartilage interface between said soft tissue andsaid bone.

In one embodiment, the scaffold apparatus comprises a nanofiber meshconfigured to apply said mechanical loading to said soft tissue graft.

In another embodiment, mechanical loading is applied by said scaffoldapparatus dynamically or intermittently to said soft tissue graft.

In another embodiment, mechanical loading is applied by said scaffoldapparatus statically to promote regeneration of a fibrocartilageinterface between said soft tissue and said bone in a subject.

In another embodiment, the scaffold apparatus comprises a material thatpromotes growth and proliferation of chondroblasts.

In another embodiment, the scaffold apparatus comprises first and secondphases, wherein (i) the first phase comprises a material that promotesgrowth and proliferation of chondroblasts, (ii) the second phaseadjacent to the first phase comprises a material that promotes growthand proliferation of osteoblasts.

In yet another embodiment, the scaffold apparatus comprises first,second and third phases, wherein (i) the first phase comprises amaterial that promotes growth and proliferation of fibroblasts, (ii) thesecond phase adjacent to the first phase comprises a material thatpromotes growth and proliferation of chondroblasts, and (iii) the thirdphase adjacent to the second phase comprises a material that promotesthe growth and proliferation of osteoblasts.

The specific embodiments and examples described herein are illustrative,and many variations can be introduced on these embodiments and exampleswithout departing from the spirit of the disclosure or from the scope ofthe appended claims. Elements and/or features of different illustrativeembodiments and/or examples may be combined with each other and/orsubstituted for each other within the scope of this disclosure andappended claims.

Further non-limiting details are described in the following ExperimentalDetails section which is set forth to aid in an understanding of thesubject matter but is not intended to, and should not be construed to,limit in any way the claims which follow thereafter.

EXPERIMENTAL DETAILS Example 1 Nanofiber Scaffolds and Human RotatorCuff Cells on Nanofiber Scaffolds

Characterization of the Aligned and Unaligned Nanofiber Scaffolds

Both aligned and unaligned PLGA nanofiber scaffolds were successfullyfabricated and characterized. Structural properties of aligned andunaligned nanofiber scaffolds are summarized in Table 1.

TABLE 1 Structural Properties of Unaligned and Aligned NanofiberScaffolds Scaffold Fiber Pore Thickness Diameter Diameter PorosityPermeability (mm) (nm) (μm) (%) (m4/N · s) Aligned 0.22 ± 0.02   615 ±152.4 228 ± 1.056 80.745 ± 2.966  (7.87 ± 2.47) × 10⁻¹² (n = 5)Unaligned 0.19 ± 0.02 568 ± 147 4.914 ± 0.777   81.760 ± 3.929 (5.72 ±0.637) × 10⁻¹² (n = 5)

The average nanofiber diameter for the aligned nanofiber scaffolds was615±152 nm, while fiber diameter of the unaligned group measured 568±147nm. No significant difference was found between the two groups.Similarly, nanofiber scaffold porosity, pore diameter and permeabilitywere also found to be comparable between the aligned and unalignednanofiber scaffolds (Table 1).

In contrast, the mechanical properties of the as-fabricated aligned andunaligned nanofiber scaffolds differed significantly (p<0.05). As shownin FIG. 2, the aligned nanofiber scaffold exhibited a markedly differentstress-strain profile when compared to the unaligned nanofiber scaffold,with a significantly higher tensile modulus, yield stress as well asultimate tensile stress. Table 2 summarizes the mechanical properties ofaligned and unaligned nanofiber scaffolds, with significantly highermechanical properties found in the aligned nanofiber scaffolds(*:p<0.05).

TABLE 2 Mechanical Properties of Aligned and Unaligned Scaffolds (*: p <0.05) Elastic Yield Strength Ultimate Modulus (Mpa) (Mpa) Stress (Mpa)Aligned (n = 5) 341 ± 30* 9.8 ± 1.1* 12.0 ± 1.5* Unaligned 107 ± 23  2.5± 0.4  3.7 ± 0.2 (n = 5)

Specifically, the aligned nanofiber scaffold exhibited a three-foldhigher elastic modulus, and nearly four-fold higher yield strength andultimate tensile strength when compared to the unaligned nanofiberscaffolds. It is emphasized here that the nanofiber scaffold structuralproperties (porosity, permeability) are similar to values reported forsoft tissue (Weiss, 2006; Yin, 2004; O'Brien, 2007; LeRoux, 2002; Joshi,1995; Kwan, 1989; Latridis, 1998; Drost, 1995) and the mechanicalproperties of the aligned nanofiber scaffolds are within range of thosereported for human supraspinatus tendon (Itoi, 1995).

Effects of Nanofiber Organization on Fibroblast Attachment and Alignment

1. Cell Attachment Morphology

The attachment morphology and growth of human rotator cuff fibroblastson aligned and unaligned nanofiber scaffolds were visualized via bothelectron and confocal microscopy. As shown in FIG. 3, the fibroblastsattached to the nanofiber scaffold but assumed distinct morphologies onthe two types of nanofiber scaffolds. Specifically, the cells grown onthe aligned fibers adopted a phenotypic elongated morphology andoriented in the direction of the long axis of the fiber. In contrast,fibroblasts seeded on the unaligned mesh exhibited a polygonalmorphology without preferential orientation. Moreover, while the humanfibroblasts proliferated on both types of substrates over time, thesemorphological differences were maintained over the two-week culturingperiod (FIG. 3).

2. Gene Expression

Fibroblast gene expression was compared over time on aligned andunaligned nanofiber scaffolds (FIG. 4). All groups expressed thehousekeeping gene GAPDH at all time points. It was observed thatintegrin expression differed between the aligned and unaligned nanofiberscaffolds while the cells expressed both types I and III collagen on thenanofiber scaffolds. The expression of α2 integrin was observed on thealigned nanofiber scaffolds at days 1, 3 and 14, while no α2 expressionwas seen on the unaligned nanofiber scaffolds. All genes evaluated wereexpressed at all time points on the monolayer control (data not shown).

3. Quantitative Analysis of Cell Attachment and Alignment

The attachment response of human rotator cuff fibroblasts to theinherent organization of the nanofiber scaffolds (aligned vs. unaligned)was further analyzed using quantitative circular statistical analysis(58, 67). Specifically, cell alignment and orientation over time werecompared to those of the underlying nanofiber substrate, focusing onmean vector angle (FIG. 5A) and angular deviation as well as mean vectorlength (FIG. 5B). Table 3 summarizes the alignment analysis results.

TABLE 3 Summary of Cell and Fiber Orientation Mean Angle ± AngularDeviation Day 1 (n = 3) (°) Aligned Cells 4.24 ± 19.73 Aligned Fibers4.32 ± 17.00 Unaligned Cells −63.46 ± 37.70  Unaligned Fibers −55.69 ±35.97  Aligned Cells 6.19 ± 18.74 Aligned Fibers 4.62 ± 17.78 UnalignedCells −19.45 ± 35.93  Unaligned Fibers −20.84 ± 38.14 

At day 1, fibroblast attachment on the aligned nanofiber scaffolds wassignificantly more aligned than on unaligned nanofiber scaffolds. Asshown in FIG. 5A, analysis of cell and fiber orientation and alignmentat day 1 revealed that cells grown on the aligned nanofiber scaffoldsare oriented horizontally, exhibiting a similar mean angle distributionand narrow angular deviation as that of the aligned nanofibers on whichthe cells are seeded. Interestingly, fibroblasts on unaligned nanofiberscaffolds also conformed to matrix organization at day 1, demonstratinga similar random orientation with a wide angular deviation approximatingthose of the unaligned nanofiber scaffold matrix (FIG. 5A, Table 3).

At day 14, fibroblast growth and morphology continued to be dictated bythe underlying nanofiber organization, with significantly higheralignment measured for fibroblasts cultured on the aligned nanofiberscaffolds compared to those found on the unaligned nanofiber scaffolds(p<0.05). However, within each nanofiber scaffold type (aligned orunaligned), circular statistical analysis of fibroblast growth revealedno significant change in alignment parameters when compared to day 1results. Specifically, the mean angle (FIG. 5A) and angular deviation(Table 3), as well as the mean vector length (FIG. 5B) values of bothcell and nanofiber scaffold samples at day 14 did not differsignificantly from those found at days 1.

Effects of Nanofiber Organization on Fibroblast Growth and MatrixDeposition

The human tendon fibroblasts proliferated on the nanofiber scaffoldsover the two week period, with no significant difference in cellproliferation found between the aligned and unaligned groups (FIG. 6A).Matrix deposition by human fibroblasts on nanofiber scaffolds was alsoevaluated over time, and the cells produced a collagen-rich matrixcontaining both type I and type III collagen (FIG. 6B). Additionally,circular statistical analysis of the immunohistochemistry imagesrevealed that collagen matrix deposition was also guided by nanofiberorganization, with an aligned type I collagen matrix found only when thefibroblasts were cultured on the aligned nanofiber scaffold (FIG. 6C).

Effects of In Vitro Culture on Nanofiber Scaffold Mechanical Properties

The mechanical properties of fibroblast-seeded PLGA nanofiber scaffoldswere also determined over time and compared as a function of in vitroculture (cellular vs. acellular nanofiber scaffolds) as well asnanofiber organization (aligned vs. unaligned). As expected for PLGA, invitro culture decreased nanofiber scaffold mechanical properties overtime, with a significantly lower ultimate tensile stress and yieldstrength found over time for both aligned and unaligned nanofiberscaffolds when compared to the as-fabricated nanofiber scaffolds (FIG.7A,C; p<0.05). In contrast, in vitro culture has no significant effecton nanofiber scaffold elastic modulus (FIG. 7B). When compared to theacellular controls, the ultimate tensile stress, elastic modulus and theyield strength of the fibroblast-seeded nanofiber scaffolds did not varysignificantly over time (FIG. 7).

At all examined time points and culture conditions, the elastic modulus,ultimate tensile stress, and yield strength of the aligned nanofiberscaffolds were significantly greater than those of the unalignednanofiber scaffolds (p<0.05), and this trend was consistently observedin both cellular and acellular groups. Interestingly, a significantdecrease in ultimate tensile stress was detected at day 1 for theunaligned nanofiber scaffold, while such a decrease was not found untila week later for the aligned group (FIG. 7A, day 7). For yield strength,a significant decrease was measured for aligned nanofiber scaffolds atday 1 compared to the as-fabricated group (FIG. 7C, p<0.05), and theyield strength of the day 14 aligned nanofiber scaffolds wassignificantly lower compared to day 1 (p<0.05). In contrast, nosignificant decrease was observed for the aligned nanofiber scaffolds atany time point.

By controlling nanofiber organization (aligned vs. unaligned), nanofiberscaffolds with controlled matrix anisotropy mimicking those of anymuscle-to-bone insertion site, e.g., a rotator cuff, may be engineered.In the present invention, the effects of nanofiber organization on theresultant nanofiber scaffold structural and mechanical properties, aswell as the response of primary cells derived from the human rotatorcuff tendons were systematically investigated. Additionally, the effectsof fibroblast culture and duration on the ability of the nanofiberscaffold to maintain stable mechanical properties were also evaluated.It was found that nanofiber organization controls nanofiber scaffoldmechanical properties and is the primary factor guiding the attachmentmorphology, alignment, gene expression as well as matrix deposition byhuman rotator cuff fibroblasts. Moreover, while in vitro cultureresulted in an expected decrease in nanofiber scaffold mechanicalproperties with polymer degradation, these changes were modulated bynanofiber organization. Based on these observations, it is clear thatnanofiber organization exerts significant control over cell response aswell as nanofiber scaffold properties, and it is a critical parameter innanofiber scaffold design for functional muscle-to-bone repair, e.g.,rotator cuff repair.

In addition to possessing a fiber diameter approximating that ofcollagen fibers present in the tendon extracellular matrix, thenanofiber scaffolds of the present invention exhibit structuralproperties that are optimal for soft tissue repair, including its over80% porosity and high permeability which can facilitate nutrienttransport as well as promote cell viability and tissue growth in vitroand in vivo. Additionally, the permeability of the nanofiber scaffold iswithin range of those reported for musculoskeletal tissues (Weiss, 2006;Yin, 2004; O'Brien, 2002; Joshi, 1995; Kwan, 1989; Latridis, 1998;Drost, 1995). For the unaligned and aligned nanofiber scaffolds of thepresent invention, tensile (elastic) modulus varied from 107 MPa to 341MPa and the ultimate tensile stress from 3.7 MPa to 12.0 MPa,respectively. While the elastic modulus of the nanofiber scaffold wasunaffected by in vitro culture, both yield strength and ultimate tensilestress decreased significantly over time. However, the magnitude ofnanofiber scaffold mechanical properties remained within the range ofthose reported for human rotator cuff tendons (Itoi, 1995). Moreover,the mechanical properties of the nanofiber scaffolds of the presentinvention can be further tailored by varying, e.g., either the averagemolecular weight or co-polymer ratio of the degradable polymers (Lu,2005).

The observed differences in mechanical properties between the alignedand unaligned nanofiber scaffolds are similar to those seen in othertypes of nanofiber scaffolds (Li, 2007; Baker, 2007; Lee, 2005;Courtney, 2006; Stankus, 2006). These reports collectively highlightanother distinct advantage of the nanofiber scaffold in that by varyingnanofiber organization and alignment, matrix anisotropy can bepre-engineered into nanofiber scaffold design. This is especiallydesirable for tendon repair or regeneration, as nanofiber scaffolds withcontrolled matrix anisotropy can be fabricated to recapitulate theinherent structure-function relationship of, e.g., rotator cuff tendons.

When Itoi et al. evaluated the tensile mechanical properties of thehuman supraspinatus tendon at the anterior, middle and posterior regionsof the tendon; it was found that both the elastic modulus and ultimatetensile stress varied as a function of tendon region. Specifically, theelastic modulus values ranged from 50 MPa to 170 MPa when progressingfrom the posterior to the superficial region, while the ultimate tensilestress spanned from 4.1 MPa to 16.5 MPa from the posterior to thesuperficial tendon region. Moreover, Thomopoulos et al. (Thomopoulos,2003) reported that collagen organization and alignment play animportant role in reducing stress concentration at the supraspinatustendon-to-bone insertion (Thomopoulos, 2006). By controlling nanofiberorganization (alignment or layering) and/or other nanofiber scaffoldparameters such as polymer composition or molecular weight, nanofiberscaffolds with biomimetic collagen alignment and region-dependentmechanical properties can be readily engineered and utilized for, e.g.,rotator cuff repair, augmentation and regeneration.

Nanofiber organization also exerted profound effects on cellularresponse. The nanofiber scaffolds of the present invention werepre-designed with similar structural properties (nanofiber diameter,porosity, pore size, permeability). Thus, any observed differences incell response can be primarily attributed to differences in nanofiberorganization (aligned vs. unaligned).

Human rotator cuff fibroblast attachment and matrix production wereguided by the organization of nanofibers of the aligned or unalignednanofiber scaffolds. Both qualitative and quantitative analyses revealedthat the cells exhibited phenotypic morphology and attachedpreferentially along the fiber axis of the aligned nanofiber scaffolds,while random cell attachment was found on the unaligned nanofiberscaffold. These differences were maintained over time and moreimportantly, subsequent cell-mediated matrix production also conformedto scaffold nanofiber organization. This contact guidance phenomena, inwhich surface topography of the biomaterial substrate regulates thespatial distribution of focal contacts and the direction of cellspreading (Singhvi, 1994; Curtis, 1997; Glass-Brudzinski, 2002;Teixeira, 2003; Wang, 2003; Engelmayer, 2006), is similar to thosereported for connective tissue cells cultured on synthetic or naturalpolymer-based nanofiber scaffolds (Li, 2007; Baker, 2007; Li, 2003; Lee,2005; Bashur, 2006; Li, 2002; Stankus, 2006; Zhong, 2006). As discussedabove, these observed differences in cell response on aligned orunaligned nanofiber scaffolds and the resultant matrix properties can bereadily exploited for, e.g., functional repair of rotator cuff injuriesas well as the formation of complex tissues.

Interestingly, the data presented herein suggest that the human rotatorcuff fibroblasts may detect differences in nanofiber alignment duringinitial attachment as well as post-adhesion matrix synthesis. It wasidentified here that while the expression of integrins such as α5 and αVwere consistently observed on both aligned and unaligned nanofiberscaffolds, α2 expression was only detectable on the aligned nanofiberscaffolds (FIG. 4). Li et al. (Li, 2003) compared fetal bovinechondrocytes response in monolayer culture to those seeded on anunaligned poly(ε-caprolactone) nanofiber scaffold, and found that α2expression was suppressed when compared to monolayer controls. It hasbeen reported that α2 is a key integrin that mediates cell attachment tocollagenous matrix (Hynes, 1992; Loeser, 2000; Hynes, 2002; Garcia,2005; Delon, 2007). Thus, expression of α2 integrin by rotator cufffibroblasts on the aligned nanofiber scaffold suggests that matrix fiberalignment may also regulate integrin expression. Moreover, compared tothe unaligned nanofibers, the aligned nanofiber scaffold may bettermimic the native extracellular matrix produced by the rotator cufffibroblasts.

Fibroblasts proliferated on both aligned and unaligned nanofiberscaffolds of the present invention over time, with no significantdifference observed between groups at all time points. These results arein agreement with previous studies which also reported minimal effect oncell proliferation due to nanofiber organization (Baker, 2007; Lee,2005). Additionally, no apparent differences in types I and III collagendeposition were observed between the aligned and unaligned groups inimmunohistochemical staining. While collagen production was notquantified in this study, Baker and Mauck (Baker, 2007) reported thatfiber architecture has little effect on normalized collagen content forboth bovine mesenchymal stem cells and meniscal fibrochondrocytescultured on aligned and unaligned poly(ε-caprolactone) nanofiberscaffolds. Interestingly, Lee et al. (Lee, 2005) reported significantlyhigher total collagen synthesis by human ligament fibroblasts grown onaligned polyurethane nanofiber scaffolds subjected to mechanicalloading. These results collectively indicate that mechanical stimulationand fiber organization may be coupled to promote overall collagenproduction over time, and in turn improve cuff healing and long termclinical response.

The intended application of the novel nanofiber scaffold systempresented here is to improve, e.g., the repair and/or augmentation oftears in muscle-to-bone insertion sites, such as, e.g., rotator cufftendons, specifically by providing a biomimetic substrate withphysiologically relevant mechanical properties that will enablefunctional and stable repair or augmentation of the damaged area, e.g.,rotator cuff. Therefore, it is important to characterize whethernanofiber scaffold mechanical properties will be maintained during invitro culture and if cell-mediated matrix production will compensate forchanges in mechanical properties due to nanofiber degradation.

As expected, nanofiber scaffold mechanical properties decreased due tohydrolytic degradation of the PLGA nanofibers (Lu, 2005). Interestingly,the aligned PLGA nanofiber scaffolds maintained the as-fabricatedtensile mechanical properties longer than the unaligned group. Theultimate tensile stress of the unaligned nanofiber scaffold decreasedsignificantly by day 1, while no such change was detected for thealigned nanofiber scaffolds until day 7. These observations also suggestthat degradation kinetics of the nanofiber scaffolds and associatedchanges in mechanical properties are dependent on fiber organization. Nosignificant increase in nanofiber scaffold mechanical properties due tofibroblast culture was observed, which is likely due to the relativelyshort culturing time evaluated, as reported significant increases inmechanical properties in cell-seeded nanofiber scaffolds are not founduntil day 70 of in vitro culture (Baker, 2007).

Effects of Co-Culture on Osteoblast and Ligament Fibroblast Phenotypes

We propose that heterotypic cellular interactions are important in themaintenance and repair of the tendon-to-bone interface. When damage tothe interface region during rotator cuff injury or subsequent repairresults in non-physiologic exposure of normally segregated tissue types(e.g., bone and tendon), heterotypic cellular interactions(osteoblast-fibroblast) may initiate repair and direct the regenerationof a neo-interface between these two tissues.

It is well documented that in post anterior cruciate ligamentreconstruction using soft tissue-based grafts, tendon-bone healingwithin the bone tunnel results in the formation of a fibrocartilage-likeinterface (Rodeo, 1993). Using a rabbit model, Koike et al. (Koike,2005) evaluated tendon-bone healing and found that after resection ofthe enthesis, the reattachment of the supraspinatus tendon to thegreater tuberosity of the humerus led to the regeneration of a newfibrocartilage-like interface. Interestingly, chondrocytes were notobserved until two weeks post reattachment and their number increased tothe level of the positive control by week six. In vitro co-culture ofligament fibroblast and osteoblast have shown that their interactionregulates cell phenotype (Wang, 2007) and results in the expression offibrocartilage-like markers such as aggrecan, cartilage oligomericmatrix protein and collagen II. These observations suggest that thefibrocartilage interface can be formed when the tendon is juxtaposedwith bone, and fibroblast-osteoblast interactions may be critical ininitiating this interface regeneration process.

Biphasic Nanofiber Scaffold for Tendon-to-Bone Interface TissueEngineering

Nanofiber scaffolds represent promising matrices for interface tissueengineering due to their superior biomimetic potential and physiologicalrelevance because they exhibit high aspect ratio, surface area, porosityand closely mimic the extracellular matrix (Ma, 2005; Christenson, 2007;Pham, 2006; Li, 2007; Murugan, 2007). These nanofiber scaffolds havebeen investigated for bone (Garreta, 2007; Fujihara, 2005; Badami, 2006;Yoshimoto, 2003), meniscus (Baker, 2007), intervertebral disc (Nerukar,2007), cartilage (Li, 2003; Li, 2005), ligament (Lee, 2005; Bashur,2006) as well as tendon tissue engineering (Sahoo, 2006).

Inspired by the structure and organization of the tendon-bone insertionsite, and focusing on mimicking collagen alignment and exercisingspatial control in mineral content, we have developed a nanofiber-basedbiphasic nanofiber scaffold. In this biphasic design, Phase A consistsof nanofibers of biodegradable polymer, such as, e.g.,polylactide-co-glycolide (PLGA) (representing the non-calcifiedinterface), while Phase B is of composite nanofibers of a biocompatibleceramic, such as, e.g., hydroxyapatite (HA) nanoparticles and abiodegradable polymer, such as, e.g., PLGA (FIGS. 10 and 17)(representing the calcified interface). It is expected that this novelnanofiber scaffold will provide the necessary structural and mechanicalproperties as well as mineral distribution to guide the regeneration ofthe complex heterogeneous tendon-to-bone interface. Additionally, byimplementing nanofibers that mimic the alignment of collagen fibers atthe insertion site, the biphasic nanofiber scaffold provides thefoundation for guided matrix deposition and tendon-bone healing.Moreover, the biphasic nanofiber scaffold exhibits tensile mechanicalproperties similar to those reported for human supraspinatus tendon(Itoi, 1995), and supports, e.g., fibroblast, osteoblast culture inpreliminary studies. The nanofiber scaffolds of the present inventionmay be fabricated into a multiphasic, e.g., a biphasic patch for, e.g.,rotator cuff repair (FIG. 10). The nanofiber scaffold can be suturedonto the tendon during cuff repair and will integrate with bone throughthe PLGA-HA phase.

Fibroblast Response on PLGA-HA Nanofiber Scaffolds

A preliminary evaluation was carried out of the mineralization potentialof human rotator cuff tendon fibroblasts as a function of HA content (0,1, 5, 15 wt %) in Phase B. Briefly, fibroblasts derived from explantcultures of human tissue (male, aged 49-79 yrs) were seeded on a PLGA-HAnanofiber scaffold according to the present invention (3.14×10⁴cells/cm²). It was found that fibroblasts remained viable (FIG. 14) andproliferated on all substrates. The cells were elongated and alignedalong the long axis of the fibers. Gene expression for alkalinephosphatase (ALP) was similar among all groups at both days 3 and 21,and only basal ALP activity levels were measured in these cultures (FIG.13). Collagen I & III deposition was maintained on the PLGA-HA nanofiberscaffolds, with no observable differences found between groups (FIG.14). These results suggest that increased HA content had no adverseeffect on fibroblast phenotype and does not appear to induce ectopicmineralization.

Co-Culture on the Biphasic Nanofiber Scaffold (Phases A & B)

To demonstrate the feasibility of co-culture on the biphasic nanofiberscaffold, bovine osteoblasts and fibroblasts derived from explantculture were labeled with Vybrant dyes (green:D\O=fibroblasts,red:D\\=osteoblasts). The fibroblasts were then seeded on Phase A (PLGA)and osteoblasts on Phase B (PLGA-HA, 5%), and allowed to attach for 15minutes before adding media. Cell distribution on each nanofiberscaffold phase, as well as the cross-section, were imaged at day 1 usingfluorescence confocal microscopy. As seen in FIG. 15, with the biphasicdesign, fibroblasts remained in Phase A while osteoblasts were localizedto Phase B. These results demonstrate the feasibility of co-culture andsuggest that the phase-specific cell distribution may lead tomulti-tissue formation.

Disclosed herein is the design and systematic in vitro evaluation of anovel biomimetic, biodegradable nanofiber scaffold for soft tissuerepair, augmentation, or replacement, such as, e.g., for functionalrotator cuff repair. The present invention discloses that nanofiberorganization has a significant effect on human rotator cuff fibroblastresponse, with the structural anisotropy of the aligned and isotropy ofthe unaligned nanofiber scaffold directly guiding cell attachment andmatrix deposition. Controlled cell response resulted in a morephysiologically relevant matrix for, e.g., rotator cuff repair on thebiomimetic nanofiber scaffold. Moreover, physiologically relevantnanofiber scaffold mechanical properties were maintained in vitro. Ourresults demonstrate that the novel nanofiber scaffold has significantpotential for enabling tendon regeneration and offers a functionaltissue engineering solution for soft tissue repair, augmentation, orreplacement, such as, e.g., rotator cuff repair.

Example 1.1 Nanofiber Scaffold Fabrication

Poly(D, L-lactide-co-glycolide) co-polymer (85:15 PLGA, Mw>>123.6 kDa;Lakeshore Biomaterials, Birmingham, Ala.) nanofiber scaffolds wereproduced via electrospinning (Matthews, 2002; Formhals, 1934; Reneker,1996). Briefly, a 35% (v/v) solution of PLGA was prepared in an organicsolvent mixture consisting of 55% N.N-dimethylformamide (Sigma-Aldrich,St. Louis, Mo.) and 10% ethyl alcohol. The polymer solution was loadedin a 5 mL syringe with a 18.5-G stainless steel blunt tip needle andelectrospun at 8-10 kV using a custom designed electrospinning device.Both aligned and unaligned nanofiber scaffolds were fabricated. Forunaligned nanofiber scaffolds, the collecting surface consisted of astationary plate, while a rotating mandrel having a diameter of 2 inchesand a length of 20 inches, which mandrel rotated at 20 m/s was utilizedto produce aligned nanofiber scaffolds. The polymer solution wasdeposited using a syringe pump (Harvard Apparatus, Holliston, Mass.; 1mL/hour) with the distance between the needle and the collecting targetdistance (air gap distance) set at 10.5 cm. See, FIGS. 19 and 20. Asshown in FIG. 20, the drum surface velocity may be varied by changingthe gear in the pump, which provides control over fiber orientation andalignment.

Example 1.2 Nanofiber Scaffold Characterization

The structural and material properties of the nanofiber scaffolds werecharacterized post fabrication. Specifically, nanofiber morphology anddiameter were imaged by Scanning Electron Microscopy (SEM, 5 kV, FEIQuanta 600, FEI Co. Hillsboro, Oreg.). The nanofiber scaffolds weresputter coated with palladium prior to SEM analysis in order to reducecharging effects. Fiber diameter was quantified by image analysis of SEMmicrographs (n=3, 2000×) using NIH ImageJ (version 1.34 s, Bethesda,Md.). In addition, nanofiber scaffold porosity and pore diameter (n=5)were determined by mercury porosimetry (Micromeritics Autopore III,Norcross, Ga.) following published protocols (Lu, 2003). In this method,the construct porosity was determined by measuring the volume of mercuryinfused into the structure during analysis. Nanofiber scaffoldpermeability (n=5) was directly determined using a custom designeddevice (51 Albro, 2006; Weiss, 2006), by first measuring the pressuredifference and then calculating permeability via Darcy's Law:K=Qh/AΔP

AΔP where k is nanofiber scaffold permeability (m⁴/N s), Q is the fluidflow rate through nanofiber scaffold (300 mL/hr), ΔP is the pressuredifference (N/m), h is the thickness of nanofiber scaffold (m) and A isthe nanofiber scaffold surface area (m²).

The mechanical properties of the as-fabricated aligned and unalignednanofiber scaffolds were evaluated under uniaxial tensile testing (Li,2002). Briefly, the nanofiber scaffolds (6 cm×1 cm) were secured withcustom clamps and mounted on an lnstron (Model 8841, Norwood, Mass.)with an average sample gauge length of 3 cm. The samples were tested tofailure at a strain rate of 5 mm/min, and the aligned nanofiberscaffolds were tested along the long axis of the aligned fibers. Boththe yield stress and ultimate tensile stress were determined, andnanofiber scaffold elastic modulus was calculated from the linear regionof the stress-strain curve.

Example 1.3 In Vitro Culture of Human Tendon Fibroblasts on NanofiberScaffolds

Cells and Cell Culture

Human rotator cuff fibroblast-like cells were obtained from explantcultures of tissue samples obtained as surgical waste following rotatorcuff repair surgery (exempted from IRB approval). For this study, thecells were derived exclusively from female patients (n=3, aged 65 to 70years). The tissue samples were maintained in Dulbecco's modified Eaglemedium (DMEM) supplemented with 10% fetal bovine serum, 1% non-essentialamino acids, 1% penicillin/streptomycin and 1% amphotericin B. Onlycells obtained from the second and third migration were used in order toensure a relatively homogeneous cell population (54). All media andsupplements were purchased from Mediatech (Herndon, Va.).

Cell Seeding on Nanofiber Scaffolds

Prior to cell seeding and to prevent nanofiber scaffold contraction(Spalazzi, 2008), the ends of aligned and unaligned nanofiber scaffoldswere secured to nanofiber scaffold holders using a sterile adhesive(Fisher Scientific, Pittsburgh, Pa.). The nanofiber scaffolds weresterilized by UV irradiation (30 minutes/side) and to promote celladhesion, the nanofiber scaffolds were pre-incubated in fullysupplemented media at 37° C. and 5% CO₂ for 16 hours. Human rotator cufffibroblasts were seeded on the nanofiber scaffolds (1 cm×1 cm cellseeding area) at a density of 3×10′ cells/cm². The cells were allowed toattach on the nanofiber scaffolds for 15 minutes, after which fullysupplemented media was added to each culture well. Cells were culturedon the aligned and unaligned nanofiber scaffolds for two weeks, and theeffects of nanofiber organization on cell morphology, attachment,proliferation and matrix production were determined at days 1, 7 and 14.In addition, the effects of in vitro culture on nanofiber scaffoldmechanical properties were also determined over the two week period.Both monolayer culture of the human tendon fibroblasts and acellularnanofiber scaffolds (aligned as well as unaligned) served as controls.

Cell Viability and Attachment Morphology

Cell attachment morphology on the nanofiber scaffolds (n=3/group) wereevaluated by SEM (FEI Quanta 600, FEI Co. Hillsboro, Oreg.) at days 1, 7and 14. The samples were first rinsed with 0.1 M sodium cacodylatebuffer (Sigma-Aldrich) and fixed in Karnovsky's fixative (Karsenty,1965; Langley, 1999) for 24 hours at 4° C., and subsequently dehydratedwith an ethanol series. The nanofiber scaffolds were coated withpalladium prior to SEM analysis to reduce charging effects. Cellviability as well as attachment morphology were evaluated by Live/Deadstaining (Molecular Probes, Eugene, Oreg.) imaged using confocalmicroscopy. Specifically, the samples (n=3/group) were rinsed twice withPBS and stained following the manufacturers suggested protocol. Thesamples were then imaged with a laser scanning confocal microscope(Olympus Fluoview IX70, Center Valley, Pa.) at wavelengths of 488 nm and568 nm.

Gene Expression

Gene expression was measured by reverse transcriptase polymerase chainreaction (RT-PCR) at days 1, 3 and 14. The nanofiber scaffolds werefirst rinsed with PBS and total RNA was isolated using the Trizolextraction method (Invitrogen, Carlsbad, Calif.). The isolated RNA wasreverse-transcribed into complementary DNA (cDNA) using the SuperscriptFirst-Strand Synthesis System (Invitrogen), and the cDNA product wasthen amplified using recombinant Taq DNA polymerase (Invitrogen).Expression of glyceraldehydes-3-phosphate dehydrogenase (GAPDH) (GAPDHsense, 5′-GGCGATGCTGGCGCTGAGTA-3′ (SEQ ID NO:1); antisense,5′-ATCCACAGTCTTCTGGGTGG-3′ (SEQ ID NO:2)), integrin α2 (sense,5′-CAGAATTTGGAACGGGACTT-3′ (SEQ ID NO:3); antisense,5′-CAGGTAGGTCTGCTGGTTCA-3′ (SEQ ID NO:4)), integrin α5 (sense,5′-GTGGCCTTCGGTTTACAGTC-3′ (SEQ ID NO:5); antisense,5′-AATAGCACTGCCTCAGGCTT-3′ (SEQ ID NO:6)), integrin αV (sense,5′-GATGGACCAATGAACTGCAC-3′ (SEQ ID NO:7); antisense,5′-TTGGCAGACAATCTTCAAGC-3″ (SEQ ID NO:8)), collagen I (sense,5′-TGCTGGCCAACTATGCCTCT-3′ (SEQ ID NO:9); antisense, 5′TTGCACAATGCTCTGATC-3′ (SEQ ID NO:10)) and collagen III (sense,5′-CCAAACTCTATCTGAAATCC-3′ (SEQ ID NO: 11); antisense,5′-GGACTCATAGAATACAATCT-3′ (SEQ ID NO: 12)) were determined. All geneswere amplified for 30 cycles in a thermocycler (Eppendorf Mastercyclergradient, Brinkmann, Westbury, N.Y.).

Quantitative Analysis of Cell Attachment on Nanofiber Scaffolds

The effects of nanofiber organization (aligned vs. unaligned) onfibroblast attachment and alignment on the nanofiber scaffolds over timewere quantified following the methods of Costa et al. (Costa, 2003).Specifically, confocal microscopy images (1024×1024 pixel resolution,n=3) of both fibroblasts seeded on the nanofiber scaffolds and acellularnanofiber scaffolds at days 1, 7 and 14 were analyzed using circularstatistics software customized for measurement of fiber alignment(Fiber3). The circular statistics parameters determined include meanvector angle (−90°≦θ≦90°; 0° indicates horizontal orientation) whichrepresents the average fiber alignment in the matrix, mean vector length(0≦r<1) which ranges from zero for a randomly distributed sample tounity for a perfectly aligned sample, and angular deviation)(0-40.5°)which characterizes the dispersion of the non-Gaussian angledistribution of the nanofibers. Specifically, angular deviation of 0°is, in general, found in a perfectly aligned sample, while 40.5° isindicative of random distribution.

Cell Proliferation

Cell proliferation (n=5) was determined at days 1, 3, 7 and 14 bymeasuring total DNA content using the PicoGreen double-stranded DNAassay (Molecular Probes) following the manufacturers suggested protocol.At designated time points, each nanofiber scaffold was rinsed twice withPBS, then treated with 0.1% Triton X solution (Sigma-Aldrich) andhomogenized by sonication (Kontes, Vineland, N.J.) in order to removeadhered cells from the nanofiber scaffold. Sample fluorescence wasmeasured with a microplate reader (Tecan, Research Triangle Park, N.C.),at the excitation and emission wavelengths of 485 nm and 535 nm,respectively. The total number of cells in the sample was determined byconverting the amount of DNA per sample to cell number using theconversion factor of 8 pg DNA/cell (Li, 2002).

Cell Matrix Production

The elaboration of types I and III collagen (n=3/group) by fibroblastsseeded on the aligned and unaligned nanofiber scaffolds were evaluatedby immunohistochemistry at days 7 and 14. Briefly, the samples wererinsed twice with PBS₁ fixed with 10% neutral buffered formalin for 24hours at room temperature. Monoclonal antibodies for type I collagen(1:20 dilution) and type III collagen (1:100) were purchased from EMDChemicals (Calbiochem, San Diego, Calif.) and Sigma-Aldrich,respectively. Before staining for type III collagen, the samples weretreated with 1% hyaluronidase for 30 minutes at 37° C. and incubatedwith primary antibody overnight. Following a PBS wash, biotinylatedsecondary antibody and Streptavidin conjugate (LSAB2 System-HRP,DakoCytomation, Carpinteria, Calif.) were added. Positive staining withthe colorimetric substrate (AEC Substrate Chromogen, DakoCytomation) wasindicated by the formation of brown precipitates and visualized with aninverted light microscope (Ziess Axiovert 25, Zeiss, Germany). At day 7,alignment of the collagen type I produced by the human fibroblasts wasalso evaluated using the circular statistics software (Fiber3) describedabove.

Example 1.4 Mechanical Properties of the Fibroblast-Seeded NanofiberScaffolds

The effects of in vitro fibroblast culture on the mechanical propertiesof the aligned and unaligned nanofiber scaffolds were determined at days1, 7 and 14. The human rotator cuff fibroblasts were grown on both typesof nanofiber scaffolds (6 cm×1 cm) at a density of 3×10⁴ cells/cm² infully supplemented media at 37° C. and 5% CO2. For the control groups,the aligned and unaligned nanofiber scaffolds without cells (acellular)were incubated in fully supplemented DMEM and analyzed at days 1, 7 and14. At each designated time point, the samples were rinsed twice withPBS, and then tested to failure under uniaxial tension following theprotocol described for the as-fabricated nanofiber scaffolds. Theelastic modulus, yield stress and ultimate tensile stress of the samples(n=5) were determined.

Example 1.5 Statistical Analysis

Results are presented in the form of mean±standard deviation, with nequal to the number of samples analyzed. One-way analysis of variance(ANOVA) was performed to determine the effects of fiber organization ofthe as-fabricated nanofiber scaffolds on material and mechanicalproperties. Two-way ANOVA was used to determine nanofiber scaffold fiberorganization and temporal effects on cellular alignment and cellproliferation. Additionally, two-way ANOVA was performed to determinethe effects of cellularity and culture time on nanofiber scaffoldtensile mechanical properties. The Tukey-Kramer post-hoc test wasutilized for all pair-wise comparisons and statistical significance wasattained at p<0.05. All statistical analyses were performed using JMPstatistical software (SAS Institute, Cary, N.C.).

Example 1.6 Chondrocytes on the Nanofiber-Based Scaffold

One objective of this experiment is to evaluate the response ofchondrocytes on PLGA and PLGA-HA nanofiber-based scaffolds. It has beenthat chondrocyte-like cells were observed at the interface during invivo tendon-to-bone healing. (Sano et al., 2002; Koike et al., 2005,2006). The second objective of this experiment is to identify an optimalnano-HA concentration for the formation of a calcifiedfibrocartilage-like matrix. It is hypothesized that chondrocytescultured on the PLGA nanofiber scaffold would form a fibrocartilage-likeECM and that scaffold mineral content would regulate chondrocyteresponse.

Experimental Procedure

Aligned PLGA and PLGA-HA nanofiber scaffolds are produced byelectrospining as discussed supra (Chun, 1995, Reneker & Chun, 1996;Moffat, et al., 2007) (PLGA (85:15)+dimethylformamide+ethanol (35%PLGA); PLGA-HA (10% and 15% HA nanoparticles, 100-150 nm, Nanocerox).

The Scaffold characterized according to the Table 4:

TABLE 4 Scaffold Characterization Measures Parameter Method ofMeasurement Surface and Bulk Properties Surface roughness: Atomic ForceMicroscopy (AFM, contact mode, pyramidal tip, k = 0.03 N/m, n = 9) FiberDiameter: Scanning Electron Microscopy (SEM, 2 kV, 4000x) + imageanalysis (Image J, n = 30) Mineral Content: Thermogravimetric analysis(TGA) (25° C.-600° C., 5° C./min, n = 5) Mineral Chemistry: FourierTransform Infrared Spectroscopy (FTIR) (200 scans, 4 cm⁻¹, n = 2)Elemental Composition: Energy Dispersive X-ray Analysis (EDAX) (10 kV,4000x, n = 2) Mechanical Properties Tension: Instron, displacement rate= 5 mm/min, n = 5 Compression: Instron, displacement rate = 2 μm/s, tareload = 10 g, n = 5

The Experiment is set up according to FIG. 72 and as follows:

The Control Group: 0% HA scaffold (PLGA); Experimental Groups: 10% and15% HA Scaffols (PLGA-HA) Medium: 10% ITS+premix, 10% FBS, 40 μg/mLL-proline, 100 μg/mL sodium pyruvate, 0.1 μM dexamethasone, 50 μg/mLascorbi acid, 1% antifungal.

Scaffold Characterization—Mineral Content

TGA confirmed the amount of minteral incorporated into the scaffold.Significantly higher mineral content was found for the 15% HA group ascompared to the 10% HA scaffolds (*p<0.05) (Table 5, FIG. 73)

TABLE 5 Scaffold Characterization - Mineral Content ThermogravimetricAnalysis (TGA) Scaffold Weight % (n = 5)  0% HA 2.99 ± 0.32 10% HA 12.69± 0.76* 15% HA 17.91 ± 0.63*

Scaffold Characterization—Composition and Chemistry

The results indicate that Ca and P peak intensity increased with mineralcontent (FIGS. 11A-11C)

Mineral chemistry of incorporated HA is confirmed by FTIR: The resultsshow phosphate bending vibration peaks at 605 cm⁻¹ and 560 cm⁻¹. This isindicative of crystalline calcium phosphate. (FIG. 74)

Scaffold Characterization—Surface Roughness and Fiber Diameter

AFM analysis revealed an increase in fiber surface roughness withmineral content (FIGS. 75A-D)

Scaffold Characterization—Fiber Diameter

Nanofiber diameter decreased with increasing mineral content (FIGS.76A-D)

Scaffold Characterization—Scaffold Mechanical Properties

Tensile modulus decreased with increasing mineral content. Significantdecrease for 15% HA case was observed. This finding is in range ofvalues reported for modulus of human supraspinatus tendon (Itoi et al.,1995) (FIG. 77A, *p<0.05).

Compressive modulus increased with increasing mineral content. There wasa significant increase for 15% HA case. This finding is in range ofvalues reported for modulus of direct insertion sites (Moffat et al.,PNAS, 2008) (FIG. 77B, *p<0.05).

Cell Morphology and Growth

Cell morphology is heterogenous with both elongated and spherical cellsobserved (FIG. 78A)

There is also significant cell proliferation within 1 week of culturewith the highest cell number on 0% HA (PLGA) scaffolds at day 42. (FIG.78B)

Chondrocyte Collagen Deposition

Collagen deposition increased over time on PLGA-HA and is significantlyhigher on the PLGA-HA. (FIG. 79A)

Dense collagen matrix is observed at day 14 for all groups with enhancedmatrix infiltration on PLGA-HA scaffolds. (FIG. 79B)

Regional distribution of collagen is found at day 42.

Nanofiber-Guided Collagen Alignment

Nanofiber scaffold organization guided collagen production with alignedcollagen matrix distributed throughout the scaffolds.

Enhanced collagen alignment was found on the PLGA scaffold as comparedto the PLGA-HA groups. (FIG. 80)

Collagen Deposition

Nanofiber scaffold supported collagen I and II production. In addition,collagen deposition is evident throughout the entire scaffold (FIG. 81).

Chondrocyte GAG Deposition

GAG synthesis increased over time for all groups and is significantlyhigher on PLGA-HA by day 28. Enhanced GAG matrix deposition andinfiltration is also observed on the PLGA-HA. (FIG. 82A-B).

Calcified Matrix Deposition

Calcified, protein-rich matrix detected on the PLGA-HA as nanofibersdegraded over time. High intensity sulfur peak associated with proteindeposition. Ca—P peak intensity ratio differed between the 10% and the15% HA groups. (FIG. 83)

Chondrocyte Hypertrophy and Matrix Maturation

Significantly higher MMP-13, Ihh and Runx2 expression was found onPLGA-HA (15%) at day 3. (FIG. 84A)

Matrix heterogeneity was also observed: nanofiber degradation,collagen-rich at periphery and along border of scaffold, GAG-rich atcenter of matrix, and uniform mineral distribution. (FIG. 84B)

Discussion

It was found that BOTH the PLGA and PLGA-HA supported: chondrocyteattachment and proliferation, synthesis of a collagen- and GAG-richmatrix, deposition of both types I and II collagen and matrix depositionguided by nanofiber alignment. These findings are indicative of theproduction of a fibrocartilage-like matrix mimicking that of the nativetendon-bone insertion.

It was also found that nanofiber compressive mechanical propertiesincreased with addition of HA. A similar trend was observed for thenon-calcified and calcified regions of the native insertion. (Moffat etal., PNAS, 2008)

In addition, designed gradient of mechanical properties exhibited byPLGA and PLGA-HA nanofiber scaffolds were found to have minimized theformation of stress concentrations and mediated load transfer betweentendon and bone.

HA content was found to affect characterization of PLGA-HA nanofibers.Similar to reported chondrocyte studies with mineral (Kandel et al.,2006; Boskey et al., 1984, Boskey, 1992; Jubeck et al., 2008), HA wasfound to increase proteoglycan and collagen production. Enhanced matrixinfiltration was also found on 15% HA.

Further, up-regulation of hypertrophic markers was found on the 15% HA.Finally, compressive mechanical properties of 15% HA similar to those ofthe native calcified fibrocartilage was found. (Moffat, et al., PNAS,2008)

Overall, these results demonstrate the potential for: 1)chondrocyte-mediated formation of the non-calcified fibrocartilageinterface region on PLGA and 2) chondrocyte-mediated formation of thecalcified interface region on the PLGA-HA (15% HA).

Example 2 PLGA-BG Composite Scaffold and Co-Cultures of Different CellTypes

To address the challenge of graft fixation to subchondral bone, a normaland functional interface may be engineered between the ligament andbone. This interface, according to one exemplary embodiment, wasdeveloped from the co-culture of osteoblasts and ligament fibroblasts ona multi-phased scaffold system with a gradient of structural andfunctional properties mimicking those of the native insertion zones toresult in the formation of a fibrocartilage-like interfacial zone on thescaffold. Variations in mineral content from the ligament proper to thesubchondral bone were examined to identify design parameters significantin the development of the multi-phased scaffold. Mineral content (Ca—Pdistribution, Ca/P ratio) across the tissue-bone interface wascharacterized. A multi-phased scaffold with a biomimetic compositionalvariation of Ca—P was developed and effects of osteoblast-ligamentfibroblast co-culture on the development of interfacial zone specificmarkers (proteoglycan, types Il and X collagen) on the scaffold wereexamined.

The insertion sites of bovine ACL to bone (see FIGS. 23A-23C) wereexamined by SEM. Pre-skinned bovine tibial-femoral joints were obtained.The intact ACL and attached insertion sites were excised with a scalpeland transferred to 60 mm tissue culturing dishes filled with Dulbecco'sModified Eagle Medium (DMEM) (see FIGS. 24A and 24B). After isolation,the samples were fixed in neutral formalin overnight, and imaged byenvironmental SEM (FEI Quanta Environmental SEM) at an incident energyof 15 keV. ACL attachment to the femur exhibited an abrupt insertion ofthe collagen bundle into the cartilage/subchondral bone matrix.Examination of collagen bundle revealed that the surface was ruffled andsmall collagen fibrils can be seen. When a cross section was imaged,three distinct zones at the insertion site were evident: ligament (L),fibrocartilage (FC), and subchondral bone (B). The interface regionspans proximally 200 μm. These cross section views showed the insertionof Sharpey fiber into the fibrocartilage (see FIGS. 25A-25C).Mineralized fibrocartilage was not distinguishable with regularcartilage from these images.

The insertion sites of bovine ACL to bone were examined by scanningelectron microscopy (SEM). Bovine tibial-femoral joints were obtained.The intact ACL and attached insertion sites were excised with a scalpeland transferred to 60 mm tissue culturing dishes filled with Dulbecco'sModified Eagle Medium (DMEM). After isolation, the samples were fixed inneutral formalin overnight, and imaged by environmental SEM (FEI QuantaEnvironmental SEM) at 15 keV.

ACL attachment to the femur exhibited an abrupt insertion of thecollagen bundle into subchondral bone. When a cross section was imaged(see FIGS. 26A and 26B), three distinct zones at the insertion site wereevident: ligament (L), fibrocartilage (FC), and subchondral bone (B).Sharpey fiber insertion into the fibrocartilage (see FIG. 8A) wasobserved. The bovine interface region spans proximally 600 μm.Examination of the interface using energy dispersive X-ray analysis(EDAX, FEI Company) enable the mineralized and non-mineralized FC zonesto be distinguished. A zonal difference in Ca and P content was measuredbetween the ligament proper and the ACL-femoral insertion (see Table 6).

TABLE 6 Region Analyzed Ca P Ca/P Ratio S Ligament 1.69 2.98 0.57 3.71Insertion 5.13 5.93 0.87 19.50

At the insertion zone (see FIGS. 27A and 27B), higher Ca and P peakintensities were observed, accompanied by an increase in Ca/P ratio ascompared to the ligament region. Higher sulfur content due to thepresence of sulfated proteoglycans at the FC region was also detected.The zonal difference in Ca—P content was correlated with changes instiffness across the interface. Nanoindentation measurements wereperformed using atomic force microscopy (AFM, Digital Instruments). Anincreasing apparent modulus was measured as the indentation testingposition moved from the ligament region into the transition zone (seeFIG. 28).

Ca—P distribution on polylactide-co-glycolide (50:50) and 45S5 bioactiveglass composite disc (PLAGA-BG) after incubation in a simulated bodyfluid (SBF) was evaluated using pCT (pCT 20, Scanco Medical,Bassersdorf, Switzerland) following the methods of Lin et al. The samplewas loaded into the system, scanned at 20 mm voxel resolution and anintegration time of 120 ms. FIGS. 29A and 29B compare the amount ofcalcified region (dark areas) observed on the PLAGA-BG disc as afunction of incubation time in SBF (from day 0 to day 28). Using customimage analysis software, it was determined that at day 0, themineralized region corresponded to 0.768% of the total disc (quartered)area, and at day 28, the mineralized region corresponded to 12.9% of thetotal area. Results demonstrate the Ca—P distribution on scaffoldsmeasured by pCT analysis.

The scaffold system developed for the experiments was based on a 3-Dcomposite scaffold of ceramic and biodegradable polymers. A compositesystem has been developed by combining poly-lactide-co-glycolide (PLAGA)50:50 and bioactive glass (BG) to engineer a biodegradable,three-dimensional composite (PLAGA-BG) scaffold with improved mechanicalproperties. This composite was selected as the bony phase of themulti-phased scaffold as it has unique properties suitable as a bonegraft.

A significant feature of the composite was that it was osteointegrative,i.e., able to bond to bone tissue. No such calcium phosphate layer wasdetected on PLAGA alone, and currently, osteointegration was deemed asignificant factor in facilitating the chemical fixation of abiomaterial to bone tissue. A second feature of the scaffold was thatthe addition of bioactive glass granules to the PLAGA matrix results ina structure with a higher compressive modulus than PLAGA alone.

The compressive properties of the composite approach those of trabecularbone. In addition to being bioactive, the PLAGA-BG lends greaterfunctionality in vivo compared to the PLAGA matrix alone. Moreover, thecombination of the two phases serves to neutralize both the acidicbyproducts produced during polymer degradation and the alkalinity due tothe formation of the calcium phosphate layer. The composite supports thegrowth and differentiation of human osteoblast-like cells in vitro.

The polymer-bioactive glass composite developed for the experiments wasa novel, three-dimensional, polymer-bioactive biodegradable andosteointegrative glass composite scaffold. The morphology, porosity andmechanical properties of the PLAGA-BG construct have been characterized.BG particle reinforcement of the PLAGA structure resulted in anapproximately two-fold increase in compressive modulus (p<0.05).PLAGA-BG scaffold formed a surface Ca—P layer when immersed in anelectrolyte solution (see FIG. 30A), and a surface Ca—P layer wasformed. No such layer was detected on PLAGA controls. EDXA spectraconfirmed the presence of Ca and P (see FIG. 30B) on the surface. TheCa, P peaks were not evident in the spectra of PLAGA controls.

In vitro formation of a surface Ca—P layer indicates PLAGA-BGcomposite's osteointegrative potential in vivo. The growth anddifferentiation of human osteoblast-like cells on the PLAGA-BG scaffoldswere also examined. The composite promoted osteoblast-like morphologyand stained positive for alkaline phosphatase, and promoted synthesis toa greater extent of Type I collagen synthesis than tissue culturepolystyrene controls.

The porous, interconnected network of the scaffold was maintained after3 weeks of culture (see FIG. 31). Mercury porosimetry (MicromeriticsAutopore III, Micromeritics, Norcross, Ga.) was used to quantify theporosity, average pore diameter and total surface area of the compositeconstruct. The construct porosity was determined by measuring the volumeof mercury infused into the structure during analysis. In addition, theconstruct (n=6) was tested under compression. BG particle reinforcementof the PLAGA structure resulted in approximately two-fold increase incompressive modulus (see Table 7, p<0.05).

TABLE 7 Pore Elastic Compressive Scaffold Average Diameter ModulusStrength Type Porosity (μm) (Mpa) (Mpa) PLAGA 31% 116 26.48 ± 3.47 0.53± 0.07 PLAGA-BG 43% 89 51.34 ± 6.08 0.42 ± 0.05

Porosity, pore diameter, and mechanical properties of the scaffold maybe variable as a function of microsphere diameter and BG content. Thegrowth and differentiation of human osteoblast-like cells on thePLAGA-BG scaffolds were also examined. The composite supportedosteoblast-like morphology and stained positive for alkalinephosphatase.

The porous, interconnected network of the scaffold was maintained after3 weeks of culture (see FIG. 31). The synthesis of type I collagen wasfound to be the highest on the composite, as compared to the PLAGA andtissue culture polystyrene (TCPS) controls (n=3, p<0.05) (see FIG. 32).

The effects of bovine osteoblast and fibroblast co-culture on theirindividual phenotypes were examined. The cells were isolated usingprimary explant culture. The co-culture was established by firstdividing the surfaces of each well in a multi-well plate into threeparallel sections using sterile agarose inserts. ACL cells andosteoblasts were seeded on the left and right surfaces respectively,with the middle section left empty. Cells were seeded at 50,000cells/section and left to attach for 30 minutes prior to rinsing withPBS. The agarose inserts were removed at day 7, and cell migration intothe interface was monitored. Control groups were fibroblasts alone andosteoblasts alone.

In time, both ACL fibroblasts and osteoblasts proliferated and expandedbeyond the initial seeding areas. These cells continued to grow into theinterfacial zone, and a contiguous, confluent culture was observed. Allthree cultures expressed type I collagen over time. The co-culture groupexpressed type Il collagen at day 14, while the control fibroblast didnot. Type X collagen was not expressed in these cultures, likely due tothe low concentration of β-GP used. Alizarin Red S stain intensity wasthe highest for the osteoblast control, (see FIG. 33C) followed by theco-cultured group (see FIG. 33B). Positive ALP staining was alsoobserved for osteoblast control and co-culture groups (see FIGS. 33F and33E respectively).

Scaffold of four continuous, graded layers with different sizes ofmicrospheres was formulated (see FIGS. 34A-347F). Layered inhomogeneitywas pre-designed into the scaffold. Due to differences in packingefficiency between different sizes of microspheres, the porosity of thescaffold decreases from layers of large microsphere to those consistingof small microspheres. PLAGA-BG composite microspheres were produced viathe emulsion method. Three layers of PLAGA-BG microspheres of differentdiameters (250-300, 300-355, 355-500 μm, from top to bottom) were used,shown in FIGS. 34A-34F. Microsphere layers were sintered at 70° C. for20 hours.

Image analysis confirmed that pore size increased from bottom to top ofscaffold. For the growth of ACL fibroblasts on the scaffold, anothertype of multi-phased scaffold was fabricated using a PLAGA mesh(Ethicon, N.J.) and two layers of PLAGA-BG microspheres. The layers weresintered in three stages in a Teflon mold. First the mesh was cut intosmall pieces and sintered in the mold for more than 20 hours at 55° C. Alayer of PLAGA-BG microspheres with diameter of 425-500 μm was thenadded to the mold. This layer was sintered for more than 20 hours at 75°C. The final layer consisted of PLAGA-BG microspheres with diametergreater than 300 μm. The scaffolds and three distinct regions werereadily observed (see FIGS. 35A-35C).

Kinetics of Ca—P layer formation on BG surfaces was related to changesin surface zeta potential in a simulated body fluid (SBF). The chemicaland structural changes in BG surface Ca—P layer were characterized usingFourier transform infrared spectroscopy (FTIR), SEM and energydispersive x-ray analysis (EDXA). FTIR provides information on thedegree of crystallinity (amorphous vs. crystalline) of the Ca—P layerformed (see FIG. 24A-24B) as well as the functional groups present on BGsurface (carbonated Ca—P layer versus non-carbonated, proteinadsorption, etc.). FTIR is much more surface sensitive than X-raydiffraction in detecting the Ca—P crystalline structures when thesurface layer is only several microns in thickness. SEM combined withEDXA is a powerful tool in relating elemental composition to specificsurface morphology and distributions (see FIGS. 25B and 25C). EDXAprovides a direct calculation of Ca/P ratio (Ca/P=1.67 for bone mineraland crystalline Ca—P layer) when appropriate standards are used. FTIR,SEM, and EDXA are complimentary techniques which together providequantitative data on the crystallinity, composition of and functionalgroups pertaining to the Ca—P layer.

We evaluated the effects of co-culturing on the growth and phenotypicexpression of osteoblasts and chondrocytes. Osteoblasts were seededdirectly on high density chondrocyte micromasses. Specific effects ofco-culture on the expression of chondrogenic markers were observedprimarily at the top surface interaction zone instead of within themicromass. Alcian blue staining (see FIG. 36B) revealed characteristicperi-cellular sulfated glycosaminoglycans (GAG) deposition bychondrocytes. GAG deposition was found largely within the micromass,instead of at the co-culture zone where elongated osteoblasts andchondrocytes were located. Sulfated GAG was not detected in thepredominantly osteoblast monolayer surrounding the micromass. Surfacechondrocytes may have de-differentiated due to co-culturing withosteoblasts. The expression of type I collagen was observed to bedistributed mainly on the top surface of the co-cultured mass (FIG.36C), where osteoblasts were located. Type I was also found at theprimarily osteoblastic monolayer surrounding the micromass (see FIG.36C, left). No type I collagen expression was observed in thechondrocyte-dominated center and bottom surface of the micromass. Highexpression of type Il collagen was observed within the micromass (seeFIG. 36D).

As types I and Il collagen were detected at the surface, it is possiblethat due to co-culture, chondrocytes and osteoblasts were forming anosteochondral-like interface at the surface interaction zone. AlizarinRed (ALZ) staining revealed that there was limited mineralization in theco-cultured group, while the osteoblast control stained increasinglypositive for calcium. It is likely that co-culture with chondrocytes mayhave delayed osteoblast mineralization. Preliminary PCR results (seeFIGS. 37A and 37B) showed that the 7 day co-culture group expressedtypes II and X collagen, as detected by RT-PCR.

Summary of Results

Effects of media additives on the growth and mineralization ofosteoblasts and human ACL fibroblasts (hACL) were examined. Duringmineralization, ALP reacted with β-glycerophosphate (βGP) and thephosphate product was utilized for mineralization. Concentrations (0,1.0, 3.0, 5.0 mM) effects were examined over time. No significant changein cell number was observed for the concentration levels of βGPinvestigated. At 1.0 mM, a significant difference between 1-day and7-day samples (p<0.05) was observed. No differences were found between1.0 mM and 3.0 mM cultures. ALZ stains for the osteoblast cultures weremore intense for 3.0 mM than for 1.0 mM. Ectopic mineralization wasobserved for hACL cultures at 3.0 mM suggesting a potential change incell phenotype.

Interaction of osteoblasts and chondrocytes on a 3-D composite scaffoldduring co-culture was examined. Scaffolds seeded with only osteoblastsor chondrocytes at the same densities served as controls. Bothshort-term and long-term co-culture experiments were conducted.Extensive SEM analysis revealed that significant interactions occurredbetween osteoblasts and chondrocytes during co-culture. Differences incellular attachment were observed between the chondrocyte controlscaffolds and the co-cultured scaffolds. On the co-cultured scaffolds,focal adhesions were evident between the spherical chondrocytes and thesurface, indicated by the arrow in FIG. 38B.

No comparable focal adhesions were observed on the chondrocyte controlsat the same time point. Chondrocyte morphology changed over time as itassumed a spherical morphology in the first 8 hours, and then spread onthe surface of the microspheres (see FIG. 33C). The nodules on thesurface of the microspheres correspond to the flattened chondrocytes.These nodules were likely chondrocytes instead of calcium phosphatenodules, since calcium phosphate nodules were approximately 1-5 μm indiameter at the culture duration observed and these nodules were about10 μm, approximately the diameter of an ovoid cell. After 7 days ofculture, the co-culture group exhibited extensive matrix production (seeFIG. 33E) and expansion on the scaffold.

Examination of the ACL-bone interface confirmed existence of a mineralgradient across the insertion zone and correlation to changes inmaterial properties. Multi-phased scaffolds with controlled morphologyand porosity were fabricated. The osteochondral graft developed fromco-culture on PLAGA-BG and hydrogel scaffold supported growth ofmultiple matrix zones with varied GAG and mineral content. BMSCsdifferentiated into ligament fibroblast and produced a functionalextracellular matrix when cultured with growth factors on a fiber-basedscaffold. Mineral content, distribution, and chemistry at the interfaceand on the scaffold were quantifiable using a complimentary set ofsurface analysis techniques (FTIR, SEM, EDAX, μCT). Electron microscopyexamination of the ACL-bone interface revealed insertion zone includingthree different regions: ligament, fibrocartilage-like zone, and bone.Co-culture of osteoblasts and ligament fibroblasts on 2-D and 3-Dscaffolds resulted in changes in cell morphology and phenotype. Type Xcollagen, an interfacial zone marker, was expressed during co-culture.Multi-phased scaffold with layered morphology and inhomogenousproperties were designed and fabricated. FTIR, SEM and EDXA arecomplimentary techniques which collectively provided qualitative andquantitative information on the Ca—P layer and composition of thecalcium phosphate surface.

Example 3 Biomimetic, Inhomogenous Triphasic Scaffold andOsteoblast-Fibroblast Co-Culture on the Scaffold

A multi-phased scaffold system with inhomogenous properties (FIG. 39)was designed and evaluated for its ability to support the growth anddifferentiation of multiple cell types. Effects of osteoblast-ligamentfibroblast co-culture on a development of interfacial zone specificmarkers (proteoglycan, types Il and X collagen) on the scaffold wereexamined.

The contiguous scaffold included three sequential phases (A-C), withPhase A (polymer fiber mesh with no Ca—P) intended for ligament culture,and Phase C (polymer-ceramic composite with high Ca—P) for boneformation. Phase 9 (polymer-ceramic composite, lower Ca—P than Phase C),the intermediate region, was where an interfacial zone developed due tothe interaction of these two cell types. The scaffolds were fabricatedfrom PLAGA 50:50, and the same polymer was used throughout. The threephases were sintered together past a polymer glass transitiontemperature to form a multi-phased scaffold. The aspect ratio betweenthe phases of the sintered cylindrical scaffold was as follows:A:B:C=2:1:2, and the as-made, complete construct was 1.0 cm in lengthand 0.40 cm in diameter (see FIG. 39).

The mineral gradient was created by incorporating differentconcentrations of bioactive glass (BG) particles during the microspheresynthesis process. BG weight percentage was correlated to the Ca—Pcontent of the interface by comparative EDXA analysis of the Ca—Psurface developed through immersion in a simulated body fluid followinga well-characterized method to create Ca—P layer on bioactive glasssurfaces as described by Lu et al. (2000) and incorporated by referenceherein.

When a specific BG weight percentage was correlated with the Ca—Pdistribution and Ca/P ratio of either the bone or the cartilage regionas described above, scaffolds were fabricated based on this weightpercentage.

The three phases of the scaffold were inhomogeneous in properties, withzonal differences in mineral content and matrix morphology (see Table8).

TABLE 8 Phase A Phase B (Ligament) (Interface) Phase C (Bone)Composition PLGA 50:50 PLGA 50:50/BG PLAGA 50:50/BG (no BG (lower)(Higher) Porosity/Pore 40% 40% 40% Diameter 100 μm 100 μm 100 μm MatrixFiber Mesh Microsphere Microsphere Morphology Based Based

The differences mimic the ACL-bone interface and facilitate the growthof different tissues. Phase C has a high mineral content compared toPhase A. While the three phases share the same polymer composition, theydiffer in weight percentage of BG. A positive correlation exists betweenscaffold stiffness and mineral content of the phase.

The three phases also differ in morphology, with Phase A composed of aporous fibrous mesh, and Phases B and C made of microsphere-based porousscaffold. Post-fabrication characterization of the scaffold includedporosity, average pore size, total surface area, as well as mechanicalproperties under compression. Scaffold porosity was held constant at 40%with a pore diameter of 100 μm, with focus on the effect of mineralcontent on cellular response as a more relevant parameter in controllingfibroblast phenotype or dedifferentiation into chondrocytes. Growth anddifferentiation of osteoblasts and ligament fibroblasts co-cultured onthe scaffold were examined. Osteoblasts were seeded on Phase C whileligament fibroblasts were seeded on Phase A.

The growth and differentiation of cells on the scaffold was monitored asa function of culturing time (1, 3, 7, 14, 21 days). Cell proliferation,ligament phenotypic expression (fibronectin, type I, III, I1 collagensynthesis, laminin, fibronectin) and osteoblast phenotype (alkalinephosphatase, type I collagen, osteocalcin, mineralization) wereexamined. Expression of interface-specific markers such asproteoglycans, types Il and X collagen were determined to assess changesin fibroblast phenotype.

The three phases of the scaffold differed in composition and morphology,while the same porosity and pore diameter were maintained. Focus wasplaced on the mineral content of the scaffold for two reasons: 1) it isa more relevant parameter for consideration of the varied mineraldistribution within the ACL-bone interface; and 2) mineral content wasutilized to direct fibroblast phenotype change or dedifferentiation intochondrocytes.

A component of the polymer ceramic composite scaffold was polylactide(PLA) which degrades via hydrolysis into lactic acid, which maycontribute to changes in ligament fibroblast phenotype. Increasedmineralization by ligament fibroblasts was observed with increasingconcentration of β-glycerophosphate, a media additive commonly used inosteoblast cultures.

The effects of co-culture were evaluated in conjunction with scaffoldmineral content. Multiple cell types were considered because theinsertion site was made up of four zones, each dominated by a specificcell type. Cell to cell interactions played a significant role indictating the formation of the interface between ligament and bone.Examination of osteoblast and ligament fibroblast co-culturesestablished that both cell types proliferated and expanded beyond theinitial seeding areas, and that a contiguous and confluent culture wasobserved at the interface after two weeks. Preliminary studies revealedthat co-culture and/or interactions with chondrocytes may have delayedosteoblast-mediated mineralization. Type X collagen was found in theosteoblast-chondrocyte co-cultured samples.

Example 4 Development of a Biodegradable, Porous, Polymer BioactiveGlass Composite Possessing Improved Mechanical Properties andOsteointegrative Potential

An objective of the experiments (described below) was to develop athree-dimensional (3-D), porous composite of polylactide-co-glycolide(PLAGA) and 45S5 bioactive glass (BG) that is biodegradable, bioactive,and suitable as a scaffold for bone tissue engineering (PLAGA-BGcomposite). Additional objectives of the study were to examine themechanical properties of a PLAGA-BG matrix, evaluate the response ofhuman osteoblast-like cells to the PLAGA-BG composite, and evaluate theability of the composite to form a surface calcium phosphate layer invitro. Structural and mechanical properties of PLAGA-BG were measured,and the formation of a surface calcium phosphate layer was evaluated bysurface analysis methods. The growth and differentiation of humanosteoblast-like cells on PLAGA-BG were also examined. The addition ofbioactive glass granules to the PLAGA matrix resulted in a structurewith higher compressive modulus than PLAGA alone. Moreover, the PLAGA-BAcomposite was found to be a bioactive material, as it formed surfacecalcium phosphate deposits in a simulated body fluid (SBF), and in thepresence of cells and serum proteins. The composite supportedosteoblast-like morphology, stained positively for alkaline phosphatase,and supported higher levels of Type I collagen synthesis than tissueculture polystyrene controls. A biodegradable, porous, polymer bioactiveglass composite possessing improved mechanical properties andosteointegrative potential compared to biodegradable polymers ofpoly(lactic acid-glycolic acid) alone was successfully developed.

PLAGA-BG Composite Disc and Microsphere Fabrication and Characterization

Polylactide-co-glycolide 50:50 co-polymer (PLAGA, Mw approximately equalto 50,000, American Cyanamide, Sunnyvale, Calif.) and 45S5 bioactiveglass (BG, MO-SCI Corporation, Rolla, Mo.) granules were used tofabricate the composite (PLAGA-BG) discs and microspheres. FIG. 40 is aschematic of the synthesis process of some forms of PLAGA-BG compositeused in this study. Specifically, PLAGA-BG discs were formed through thetraditional solvent-casting process, where PLAGA and BG granules werefirst mixed according to a polymer to ceramic weight ratio of 1:3 anddissolved in methylene chloride. The solution was then slowly pouredinto a Teflon mold and allowed to cool overnight in a −20° C. freezer.The resultant polymer-ceramic film was bored into 1-cm wide and 0.1-mmthick discs. The discs were then dried overnight to remove any residualsolvent (Lyph-lock 12, PLAGA-BG composite microspheres were formedthrough a water-oil-water emulsion. Specifically, PLAGA granules werefirst dissolved in methylene chloride, and BG particles (<40 μm) wereadded to achieve a 25% mixture. The mixture was then poured into a 1%polyvinyl alcohol (Polysciences, Warrington, Pa.) solution. Thesuspension was stirred constantly, and the spheres were allowed toharden in the polyvinyl alcohol solution. The resultant microsphereswere then washed, vacuum filtered, and dried at room temperature. Next,the composite microspheres were sifted using a mechanical sifter to afinal size range of 100-200 μm. The cylindrical construct, averaging 0.5cm in width and 1.0 cm in height, was fabricated by heating themicrospheres at 70° C. for 20 hours in a stainless-steel mold.

Before in vitro evaluations, the morphology, porosity and mechanicalproperties of the PLAGA-BG construct were determined. Poreinterconnectivity, morphology, and the bonding of microspheres withinthe construct was examined by scanning electron microscopy (SEM, Amray1830-D4), at an acceleration voltage of 20 keV. Elemental composition ofthe composite surface was determined by energy-dispersive X-ray analysis(EDXA). Mercury porosimetry (Micromeritics Autopore III, Micromeritics,Norcross, Ga.) was used to measure the porosity, average pore diameter,and total surface area of the composite construct. In this method, theconstruct porosity was determined by measuring the volume of mercuryinfused into the structure during analysis. In addition, the construct(n=6) was tested under compression using the lnstron ServohydrolicSystem 8500 (Instron, Canton, Mass.), with a ramp speed of 0.02 cm/s.The compressive strength and elastic modulus of the construct weredetermined. PLAGA scaffolds without BG served as controls.

The composite discs were immersed for 1, 7, and 14 days in a simulatedbody fluid (SBF) whose ion concentration is similar to that ofextracellular fluid. PLAGA discs without BG served as controls. Asurface area to volume ratio of 1.0 cm″¹ was maintained for allimmersions. The pH of the solution as a function of immersion time wasmeasured. Perfect sink conditions were maintained during the immersionstudy. SEM (Amray 1830-D4) and EDXA were used to monitor the formationof a Ca—P layer on composite films.

Seeding Human Osteosarcoma Cells on the PLAGA-BG Composite Scaffold

Human osteosarcoma cells (SaOS-2) were cultured in Medium 199 (M199,Sigma Chemicals, St. Louis, Mo.), supplemented with 10% fetal bovineserum (Life Technologies, Rockville, Md.), L-glutamine, and antibiotics.The cells were grown to confluence at 37° C. and 5% CO₂. Under theseconditions, the osteoblastic phenotype of SaOS-2 was maintained for upto at least four weeks of culture, with positive expression of alkalinephosphatase, type I collagen, osteocalcin, and formation of mineralizedcultures.

SaOS-2 cells were seeded on the porous, PLAGA-BG scaffolds (n=3) at thedensity of 5×10⁴ cells/cm², and were cultured in 12-well plates (FisherScientific, Fair Lawns, N.J.) for up to 3 weeks. PLAGA alone and tissueculture polystyrene (TCPS) served as control groups. Once the cells havegrown to confluence, at two weeks from the start of culture,mineralization medium containing 3.0 mM of β-glycerophosphate and 10μg/ml of L-ascorbic acid

Cell adhesion and growth morphology on the 3-D construct were monitoredusing SEM (20 keV). Alkaline phosphatase staining was performed at eachculturing time point, using a standard histochemical assay. The sampleswere incubated for 30 min with Napthol AS—Bi (Sigma), phosphate salt,N,N-dimethyl formamide (Sigma), and Fast Red (Sigma) at 37° C. Thesamples were then fixed in 2% paraformaldehyde for 30 min at 4° C. Thesynthesis of type I collagen by SaOS-2 cells was quantified using amodified ELISA.

The formation of mineralized nodules was examined by SEM-EDXA.Mineralization was further ascertained using Alizarin Red S staining forcalcium. Briefly, the samples were washed with deionized H₂O, fixed with2% paraformaldehyde and incubated in 2% Alizarin Red S solution for 5min. The samples were then washed with deionized water and viewed underthe microscope.

Summary of Results:

Data in the graphs are presented in the form of mean±standard deviation(mean±SD), with n equal to the number of samples analyzed per immersiontreatment. One-way analyses of variance (ANOVA) and the Student's t-testwere used to compare the mechanical testing data (n=6), porosimetryresults (n=3), as well as the collagen synthesis data (n=3). Statisticalsignificance was evaluated at the p<0.05.

SEM examination of the PLAGA-BG discs revealed a homogenous distributionof the BG particles within the PLAGA phase. In addition, the compositesin disc form as well as microsphere form were visually more opaque thanPLAGA alone, largely because of the addition of BG.

Sintering of the microspheres resulted in a well-integrated structure,with the microspheres joined at the contact necks. SEM analysis revealedthat a 3-D, interconnected porous network was found throughout thecomposite construct. Elemental analysis using EDXA showed that thecomposite surface was largely made up of C, Na, Si, Ca, and P before anyimmersions.

The result from structural characterizations of the as-fabricatedcomposite scaffold are summarized in the following Table 9.

TABLE 9 Pore Elastic Compressive Scaffold Average Diameter ModulusStrength Type Porosity (μm) (Mpa) (Mpa) PLAGA 31% 116 26.48 ± 3.47 0.53± 0.07 PLAGA-BG 43% 89 51.34 ± 6.08 0.42 ± 0.05

BG particle-reinforcement of the PLAGA structure resulted in a neartwo-fold increase in compressive modulus. The structural and mechanicalproperties of the scaffold can be systematically optimized by varyingmicrosphere and scaffold fabrication parameters. Porosimetry analysisrevealed that the 3-D composite measured an average porosity of 43%,with a mean pore diameter of 89 μm. The PLAGA control scaffold exhibited31% total porosity and a mean pore diameter of 116 μm. The PLAGA-BGcomposite possessed a higher elastic modulus (51.336±6.080 MPa versus26.479±3.468 MPa) than the control PLAGA scaffold. Although the meanswere different, the compressive strength of the composite at 0.417±0.054MPa was not statistically different from that of the PLAGA control(0.533±0.068 MPa), at p<0.05.

The bioactivity of the composite was determined by monitoring theformation of a calcium phosphate layer on the composite discs in asimulated body fluid (SBF). The composite was found to be bioactivebecause it formed a calcium phosphate layer on its surface afterimmersion in SBF for 7 days. SEM-EDXA results showed that an amorphouscalcium-phosphate layer was found on the composite surface after 7 daysof immersion, whereas no such layer was detected on the control polymerwithout bioactive glass particles for the same duration. In particular,polymer-ceramic composite (PLAGA-BG) which were immersed in simulatedSBF for 14 days formed a surface calcium phosphate layer (Ca, P presenceconfirmed by X-ray analysis as summarized in FIG. 41). No such layer wasfound on the PLAGA control without 45S5 bioactive glass. The composite(PLAGA-BG) surface was covered with calcium phosphate nodules after 14days of immersion. In contrast, the PLAGA control surface, afterimmersion for 14 days in SBF, did not form a calcium phosphate layer,but began to exhibit surface pores formed due to the degradation of thepolymer.

FIG. 41 shows EDXA spectra of the PLAGA-BG composite immersed in a SBFfor 14 days. The composite surface still contained C, Si, Ca, and P,whereas the Cl peak was detected after immersion in SBF. A surfacecalcium phosphate layer has formed on the PLAGA-BG composite surface.The Ca and P peaks were not found in the spectra of PLAGA controls.

The microsphere-based, porous, PLAGA-BG composite supported the growthand phenotypic expression of human osteoblast-like cells. Media pHvariation was measured for the full duration (3 weeks) of cell culturewith PLAGA-BG and PLAGA, and physiological pH (7.3-7.7) was maintainedin all cultures for up to 3 weeks. There was no significant change insolution pH after 2 weeks of culture with osteoblast-like cells, andculture media was exchanged every other day to remove metabolic productsand supply fresh nutrients to the cells. Extensive cellular growth wasdetected on the scaffold surface as well as within the PLAGA-BGcomposite. In addition, the porous network of the scaffold wasmaintained even after 3 weeks of culture. In many areas, cellular growthhad bridged two or more microspheres while maintaining the porousstructure. SEM analysis revealed the synthesis of collagen-like fibersby the SaOS-2 cells. All cultures stained positively for the synthesisof alkaline phosphatase, although a much higher intensity of stain wasobserved in cultures with the PLAGA-BG scaffold than for PLAGA cultures.

As shown in FIG. 32, the synthesis of type I collagen by SaOS-2 cellsincreased with culturing time, with the highest amount found on PLAGA-BGcomposite (0.146±0.006 μg), as compared to PLAGA (0.132±0.006 μg), andTCPS controls (0.073±0.005 μg). The expression of type I collagen bySaOS-2 cells cultured on the composite was significantly higher thancells grown on TCPS controls, (p<0.05). There was a trend towards higherType I collagen synthesis on the PLAGA-BG composite compared to PLAGAalone, but this was not found to be significant. (p=0.06) The formationof a mineralized matrix was confirmed by positive staining with AlizarinRed S and elemental analysis in which Ca and P were detected on PLAGA-BGscaffolds cultured with SaOS-2 cells. Alizarin stain intensity increasedwith culturing time. The mineralized nodules were not observed on PLAGAor TCPS controls after 2 weeks of culture, before the addition of themineralization medium. After 1 week of culturing with the mineralizationmedium, mineralization as reflected in staining intensity, was much lesson the control substrates than on PLAGA-BG.

SEM and EDXA analyses confirmed the formation of calcium phosphatenodules on the composite surface after only 3 days of culture, beforethe addition of the mineralization medium. These calcium phosphatenodules are similar in size and shape as observed on PLAGA-BG discs inthe SBF. In time, the Ca—P nodules increased in size and formed largeraggregates, indicating that the PLAGABG composite was bioactive invitro. The relative Ca to P peak ratio of the deposits decreased as afunction of culturing time. These results collectively suggest that thecomposite was bioactive, and was capable of forming a surface calciumphosphate layer.

Example 5 ACL-Bone Interface Regeneration through Biomimetic ScaffoldDesign and Co-Culture of Osteoblasts and Fibroblasts

The degree of graft integration is a significant factor governingclinical success and it is believed that interface regenerationsignificantly improves the long term outcome. The approach of this setof experiments was to regenerate the ACL-bone interface throughbiomimetic scaffold design and the co-culture of osteoblasts andfibroblasts. The interface exhibits varying cellular, chemical, andmechanical properties across the tissue zones, which can be explored asscaffold design parameters. This study describes the design and testingof a multi-phased, continuous scaffold with controlled heterogeneity forthe formation of multiple tissues. The continuous scaffold consists ofthree phases: Phase A for soft tissue, Phase C for bone, and Phase B forinterface development. Each phase was designed with optimal compositionand geometry suitable for the tissue type to be regenerated. Fibroblastswere seeded on Phase A and osteoblasts were seeded on Phase C, and theinteractions of osteoblasts and fibroblasts (ACL and hamstring tendon)during co-cultures on the scaffolds were examined in vitro.

Phases A, B and C consist of poly(lactide-co-glycolide) (PLAGA, 10:90)woven mesh, PLAGA (85:15) microspheres, and PLAGA(85:15)/Bioactive Glass(45S5.BG) composite microspheres, respectively. The microspheres wereformed via a double emulsion method, and the continuous multi-phasedscaffolds were formed by sintering above the polymer glass transitiontemperature. Scaffold porosity and pore diameter were determined byporosimetry (Micromeritics, n=3) and the samples were tested underuniaxial compression (MTS 810, n=5) at 1.3 mm/min up to 5% strain with10 N preload.

Bovine and human osteoblasts (bOB and hOB), and bovine ACL fibroblasts(bFB) and human hamstring tendon fibroblasts (hFB) were obtained throughexplant culture. In experiment I, bOB and bFB (5×10⁵ cellseach/scaffold) were co-cultured on the scaffold, and cell viability,attachment, migration and growth were evaluated by electron andfluorescence microscopy. The bOB were pre-labeled with CM-DiI, and bothcell types were labeled with calcein AM (Molecular Probes) prior toimaging. Matrix production and mineralization were determined byhistology. After ascertaining cell viability on the scaffolds, a moreextensive experiment using hOB and hFB was conducted in which cellproliferation and differentiation and above analyses were investigated.The mechanical properties of the seeded scaffolds were also measured asa function of culture time.

Compression testing of scaffolds indicated an average modulus of 120±20MPa and yield strength of 2.3 MPa. The intrusion volume, porosity andpore diameter data are summarized in Table 10 below.

TABLE 10 Intrusion Mode Pore Volume (μL) Porosity (%) Diameter (μm)Phase A 41 ± 8 58 ± 5 73 ± 11 Phase B 28 ± 7 34 ± 4 75 ± 7  Phase C 12.525.7 83

The fibroblasts and osteoblasts were localized primarily at the two endsof the scaffolds after initial seeding, with few cells found in Phase B.After 28 days, both cell types migrated into Phase B (FIG. 42B), andextensive cell growth was observed in Phases A and C (FIGS. 42A and42C).

Extensive collagen-rich matrix production was found throughout the threephases at day 28 (FIGS. 42D-E).

The biomimetic, multi-phased scaffolds supported the growth and ECMproduction of both osteoblasts and fibroblasts. After 28 days ofculture, collagen production was evident in all three phases andmineralized matrix was found in the bone and interface regions.Osteoblast and fibroblast interaction at the interface (Phase B)suggests that these cells may play a significant role in the developmentof a functional insertion site. These findings demonstrate that thisnovel scaffold is capable of simultaneously supporting the growth ofmultiple cell types and can be used as a model system to regenerate thesoft tissue to bone interface. Additional studies can focus on scaffoldoptimization and the development of the interface on the novel scaffold.

Example 6 Two Triphasic Scaffolds for Interface Tissue Engineering

This set of experiments is directed to the development of a multi-phasedscaffold with controlled heterogeneity for interface tissue engineering.This continuous scaffold is comprised of three phases with Phase Adesigned for ligament formation, Phase C for bone, and Phase B forinterface development. The design objective was to formulate a scaffoldthat is able to support the growth and differentiation of bothosteoblasts and ligament fibroblasts. Two design parameters were variedamong the three phases: mineral (Ca/P) content and geometry. This studyintroduces a 3-D biomimetic substrate for interface development. Theinteraction of osteoblasts and ACL fibroblasts during co-culture on themulti-phased scaffold were examined. An objective of the study was todemonstrate that both cell types proliferate and elaborate a collagenlike matrix on the 3-D scaffolds.

Two types of scaffolds were fabricated. The first type is comprisedentirely of microspheres formed via a double emulsion method. Phase Aconsists of poly(lactide-co-glycolide) 50:50 (PLAGA), Phase C ofPLAGA/Bioactive glass (PLAGA-BG) composite microspheres, and Phase Bcontains a mixture of PLAGA and PLAGA-BG. For the second type ofscaffold which has a different geometry and degradation rate, Phase Aconsists of PLAGA (10:90) woven mesh, Phase C of PLAGA 85:15/BGmicrospheres, and Phase B contains PLAGA (85:15) microspheres. Thecontinuous multi-phased scaffolds were formed by sintering above theglass transition temperature.

Bovine osteoblasts and ACL fibroblasts were obtained from explantcultures of tissue isolated from neonatal calves. The cells werecultured in Dulbecco's Modified Eagles Medium (DMEM, Mediatech),supplemented with 10% fetal bovine serum, L-glutamine, and 1%penicillin/streptomycin (Mediatech).

Scaffolds were sterilized by ethylene oxide and fibroblasts were seededat a density of 5×10⁵ cell/scaffold onto Phase A, while osteoblasts wereseeded at 5×10⁵ cell/scaffold on Phase C. Phase B was left unseeded andthe migration of osteoblasts and fibroblasts into this interfacialregion was examined. The osteoblasts were labeled with CM-DiI celltracer (Molecular Probes), and their location was tracked with respectto fibroblasts and each phase of the scaffold. The scaffolds werecultured in supplemented DMEM for up to 28 days. Ascorbic acid (10μg/mL) and 3 mM β-glycerophosphate were added to the cultures at day 7.

Cell migration, attachment and growth were examined using scanningelectron microscopy (5 kV, JEOL 5600LV). Cell viability and migrationwere evaluated by fluorescence microscopy (Zeiss Axiovert 40) usingcalcein AM tracer (Molecular Probes). Matrix production andmineralization were determined via histology. The samples were fixed,embedded and sectioned, after which Trichrome, von Kossa and PicrosiriusRed stains were performed.

At day 0, SEM analysis showed that a large number of cells attached toPhase A and C of the scaffolds (FIG. 43A). Fluorescence microscopyrevealed that fibroblasts and osteoblasts were localized primarily atopposite ends of the scaffolds after initial seeding, with very fewcells found in Phase B (FIGS. 44A-C). At day 28, SEM analysis revealedthat both cell types elaborated extracellular matrix (ECM) on Phases Aand C (FIGS. 43B and 43C) with some matrix formation observed in Phase B(FIG. 43D). Fibroblasts were found largely in Phase A and osteoblasts inPhase C (FIGS. 44D and 44F), with a mixture of cell types found in PhaseB (FIG. 44E).

Histological analyses confirmed cell migration into Phase B and matrixproduction throughout the three phases of the scaffold at day (FIGS.49A-49C). The collagen-rich matrix (FIGS. 49D and 49E) seen in all threephases and osteoblast-mediated mineralization were observed on thesurface of the PLAGA-BG microspheres (FIG. 49F, see arrow).

The biomimetic, multi-phased scaffolds supported the growth and ECMproduction by both osteoblasts and fibroblasts. After 28 days ofculture, collagen production was evident in all three phases andmineralized matrix was found in the bone and interface regions only.

Osteoblast and fibroblast interaction at the interface (Phase B)suggests that these cells may serve a significant role in thedevelopment of a functional insertion site. The results demonstrate thatthis novel scaffold is capable of simultaneously supporting the growthof multiple matrix zones. Additional studies can examine the effects ofcell-cell interactions at the interface region and optimize the scaffoldfor clinical utilization.

Example 7 A Novel Micro-Co-Culture Model and Examination ofOsteoblasts-Fibroblasts Interaction in a Micro-Co-Culture

It is believed that fibroblasts and osteoblasts interactions play asignificant role in interface formation. In vivo, fibroblasts andosteoblasts form a fibrocartilage layer within the bone tunnel. Sincethe natural interface spans less than 400 pm, a novel micro-co-culturemodel was developed that utilizes microfluidics to exert spatial controlin cell distribution. This can be used to determine how cell-cellinteractions may regulate interface remodeling locally at themicro-scale. The fabrication parameters of this model were optimized andinitial osteoblastic and fibroblastic responses were examined.

Channels were designed having a bimodal non-intersecting serpentinegeometry with 200 μm features. The design was implemented on siliconwafers using SU-8 25 (Microchem) photoresist and a mold patterned usingPolydimethylsiloxane (PDMS, Dupont). In this design, osteoblast andfibroblast channels were first separated by PDMS, which was laterremoved to allow cell to cell interactions.

In order to optimize the channel depth for subsequent co-culturestudies, the spin-coating durations (30, 45, 60 and 90 s) were varied.Cell seeding time was optimized by incubating the cells within thechannels for 1, 3, 6, and 24 hours prior to removal of the PDMS followedby live-dead staining.

Bovine primary osteoblasts and fibroblasts were obtained from explantcultures. The cells were grown in supplemented DMEM (10% FBS, 1% NEAAand 1% antibiotics) at 37° C. and 5% CO₂. Osteoblast or fibroblastsuspension (20×10⁶ cells/ml) was perfused into its respectivemicrochannels. Cells were allowed to attach for 1 hour prior to PDMSremoval. Cell migration was tracked by labeling fibroblasts with CM-DiIand osteoblasts with CFDA-SE (Molecular Probes) prior to seeding.

Analyses were performed at days 1, 2, and 6 following PDMS removal.Alkaline Phosphatase (ALP) activity was ascertained with fast-blue stain(Sigma), while type-1 collagen deposition was examined byimmunohistochemistry.

A spin-coating duration of 30 seconds was chosen to balance channeldepth and uniformity. Based on the cell viability, the optimal cellattachment time within the channels was 1 hour (FIG. 46-2 a). Both celltypes migrated and proliferated beyond their initial seeding zone (FIGS.30-1 a through 46-1 d) and grew into physical contact by day 1 (FIGS.46-1 e and 46-1 f). Local confluency and cross-migration were observedat day 2. ALP activity was observed in the osteoblast region (FIG. 46-2b), while type-1 collagen was found in all regions (FIG. 46-2 c).

A successful micro-co-culture model was developed and initialexamination of the interactions between osteoblasts and fibroblasts in amicro-co-culturing environment was performed. Cells proliferated beyondthe initial seeding region and maintained their phenotypes as indicatedby ALP activity of osteoblasts and type-1 collagen deposition of bothcell types. The cell-to-cell cross-migration at day 2 offered a host ofhomotypic and heterotypic cell interactions. Micropatterning of cellsoffers an unique opportunity to control the local micro-environment andpermit the in-depth examination of cell-cell interactions. Thisunderstanding can aid in the identification of mechanisms drivinginterface formation.

Example 8 In Vitro Evaluations of Human Osteoblasts and FibroblastsCo-Cultured on Multi-Phased Scaffolds

This set of experiments was directed to in vitro evaluations of humanosteoblasts and fibroblasts co-cultured on multi-phased scaffolds. Aschematic of the experimental design for the in vitro study is shown inFIG. 47. Phase A (mesh) was seeded with human hamstring tendonfibroblast cell suspension. Phase C was seeded with osteoblasts. Cellinteraction in the interfacial Phase B was monitored over time.Acellular scaffolds served as controls.

Cell proliferation in Phases A, B, and C during 35 days of humanhamstring tendon fibroblast and osteoblast co-culture on multiphasedscaffolds is shown in FIG. 48A. A general trend of increasing cellnumber was observed in each phase over time. Data demonstrates that allthree phases of the scaffold support cellular viability andproliferation. A higher number of cells were seeded on phase A due toits inherently larger surface area compared to phase C.

Mechanical testing data for multiphased scaffolds seeded with humanhamstring tendon fibroblasts and human osteoblasts over 35 days ofculture (n=4) is graphically shown in FIGS. 48B and 48C. Scaffolds weretested in uniaxial compression. Compressive modulus (FIG. 48B) and yieldstrength (FIG. 48C) were calculated from the resulting stress-straincurves. Both cell seeded (C) and acellular (AC) scaffolds were examinedat days 0, 7, 21, and 35.

Compared to the acellular controls, the cell seeded scaffolds degradedslower and better maintained their structural integrity over time. Theyield strength of the acellular scaffold decreased over 35 days, whilethe seeded scaffolds maintained its yield strength.

Example 9 Triphasic Scaffold Fabrication and Evaluation of HumanHamstring Tendon Fibroblasts and Trabecular Bone Osteoblasts Seeded onthe Triphasic Scaffold

The scaffold designed for this study consisted of three phases and werefabricated in four stages (FIG. 49A). First, Phase A was formed frompolyglactin 10:90 poly(lactide-co-glycolide) (PLGA) mesh sheets (VicrylVKML, Ethicon). Mesh sheets were cut into small segments (approximately5 mm×5 mm) and inserted into cylindrical molds (7.44 mm diameter). Moldswere heated to 150° C. for 20 hours to sinter the segments together toform a cylindrical mesh scaffold. The next phase (Phase B) consisted of100% 85:15-poly(DL-lactide-co-glycolide) (PLAGA, Alkermes Medisorb,M_(W)26 123.6 kDa) microspheres formed by a water/oil/water emulsion.Briefly, 1 g PLAGA was dissolved in 10 mL methylene chloride (EMScience, Gibbstown, N.J.) and poured into a mixing 1% PVA surfactantsolution (Sigma Chemicals, St. Louis, Mo.). Microspheres were mixed for4 hours, recovered by filtration, allowed to dry in a fume hoodovernight, then vacuum desiccated for 24 hours. To form the PLAGAmicrosphere phase, about 0.075 g microspheres were inserted into thesame molds as used previously, and sintered at 55° C. for 5 hours. Thelast phase (Phase C) consisted of composite microspheres formed from an80:20 ratio of PLAGA and 45S5 bioactive glass (BG, MO-SCI Corporation,Rolla, Md.). Again, microspheres were formed by emulsion, except with0.25 g bioactive glass suspended in a solution of 1 g PLAGA in mLmethylene chloride. Microspheres (28-30 mg/scaffold) were sintered inthe same molds at 55° C. for five hours. After all three phases weresintered separately, Phases A and B were joined by methylene chloridesolvent evaporation, and then sintered to Phase C for 10 hours at 55° C.in the same molds. Subsequently, scaffolds were sterilized with ethyleneoxide. Final scaffold dimensions are detailed in FIGS. 32-4A and 32-4B.

Human osteoblast-like cells and hamstring tendon fibroblasts wereobtained from explant culture of tissue isolated from humerus trabecularbone and hamstring tendon respectively. Trabecular bone was rinsed withPBS, then cultured in Dulbecco's Modified Eagle's Medium (DMEM,Mediatech, Herndon, Va., USA) supplemented with 10% fetal bovine serum,1% non essential amino acids, and 1% penicillin/streptomycin (Mediatech,Herndon, Va.), and incubated at 37° C. in a 5% CO₂ incubator to allowfor cell migration. Hamstring tendon obtained from excess tissueutilized for hamstring tendon ACL reconstruction autografts was mincedand cultured in similarly supplemented DMEM. The first migrations ofcells were discarded to obtain a more uniform cell distribution. Secondmigration, passage 2 osteoblast-like cells and second and thirdmigration, passage 5 hamstring tendon fibroblasts were utilized for theco-culture experiment.

Scaffold dimensions were measured prior to cell seeding and before andafter EtO sterilization. Phase thickness was calculated by imageanalysis, while phase diameter was determined using a digital caliper.Scaffold porosity and pore diameter (Phases A and B: n=3; Phase C: n=1)were determined by mercury porosimetry (Micromeritics Autopore III andAutopore IV 9500, Micromeritics, Norcross, Ga.). The porosity data wereutilized to determine cell seeding densities and cell suspension volumesfor Phases A and C, with the volumes calculated such that fibroblastssuspension remains in Phase A and osteoblasts suspension in Phase C.

Hamstring tendon fibroblasts were seeded at a density of 250,000cells/scaffold in a volume of 40.7 μL/scaffold on Phase A (FIG. 49B).After allowing the fibroblasts to attach to the scaffolds for 20minutes, the scaffolds were rotated upside down so that Phase C facedupwards. Subsequently, 75,000 osteoblast-like cells were seeded perscaffold in a volume of 12.5 μl. After allowing the osteoblasts toattach to the scaffold for 20 minutes, the scaffolds were covered withDMEM supplemented with 10% FBS, 1% NEAA, and 1% penicillin/streptomycin,and incubated at 37° C. and 5% CO₂. Ascorbic acid at a concentration of20 μg/mL was added beginning at day 7. Media was exchanged every twodays. Scaffolds were cultured in 6-well plates and covered with 7 ml ofsupplemented media per scaffold to minimize pH fluctuations due to rapidpoly(glycolic acid) degradation.

Cell attachment, migration, and proliferation on the multi-phasedscaffolds were examined using SEM (5 kV, JEOL 5600LV) at days 7, 21, and35. The scaffolds were fixed with Karnovsky's glutaraldehyde fixative,and stored at 4° C. for 24 hours. The samples were then rinsed withHank's buffered salt solution two times, and serially dehydrated withethanol. Cross-sections of the scaffold phases were mounted on analuminum post and gold-coated prior to analysis.

Extracellular matrix production and mineralization were determined viahistology at day 35. Scaffolds were rinsed two times with roomtemperature PBS. The scaffolds were then covered with 10% neutralbuffered formalin and stored at 4° C. Samples were plastic embeddedusing a modification of a procedure developed by Erben. The scaffoldswere first suspended in 2% agarose (low gelling temperature, cellculture grade, Sigma, St. Louis, Mo.), then serially dehydrated withethanol and cleared with xylene substitute (Surgipath, Sub-X, Richmond,III.). Following dehydration, samples were embedded in poly(methylmethacrylate) (Polysciences, Inc., Warrington, Pa.) and sectioned into10 μm slices. The scaffold sections were stained with either hematoxylinand eosin, von Kossa or Picrosirius Red stains and imaged with lightmicroscopy.

At days 1, 7, 21, and 35, scaffolds were rinsed twice with PBS andsubsequently the three phases were separated. Each phase was then storedin 0.1% Triton-X at −80° C. Cellular proliferation in each phase wasdetermined by means of PicoGreen DNA quantitation assay. In addition,cellular phenotype for mineralization was evaluated using a quantitativealkaline phosphatase (ALP) assay.

At days 0, 7, 21, and 35, seeded and acellular scaffolds were testedunder uniaxial compression (MTS 810, n=4). The crosshead speed was 1.3mm/min, and the scaffolds were compressed up to 35-40% strain. A 10 Npreload was applied prior to testing. The effects of scaffolddegradation and extracellular matrix production on scaffold compressivemodulus were examined.

Mercury porosimetry data for each phase are summarized in Table 11 shownbelow.

TABLE 11 Intrusion Pore Area Volume Porosity Mode Pore (mm²) (μL) (%)Diameter (μm) Phase A (n = 3) 6000 + 800 41 ± 8 58 ± 5 73 ± 11 Phase B(n = 3) 2400 + 500 28 ± 7 34 ± 4 75 ± 7  Phase C (n = 1) 706.3 12.5 25.783

Scaffold dimensions are shown in FIGS. 50B and 50C. The thickness ofPhase C decreased significantly (p<0.05) due to contraction during theEtO sterilization (FIG. 50B). In addition, the thicknesses of all phaseswere significantly different from each other after sterilization.Scaffold diameters also varied due to contraction during sintering, inthe case of Phase A₁ and contraction of Phase C during sterilization.The diameters of Phases B and C decreased significantly aftersterilization, and the diameters of all phases were significantlydifferent from each other after sterilization (p<0.05). During thescaffold fabrication process, microspheres are lost between weighing andfilling the molds. This loss is mainly due to static charge accumulationin one or more of the microspheres, weighing paper, or mold, whichprevents a small percentage of the microspheres from entering the molds.PLAGA-BG microspheres for Phase C generally experience a 2.1±1.4% lossin mass, while the PLAGA microspheres for Phase B suffer a loss of4.0±1.8% (FIG. 50A). Composite microspheres are generally morestatically charged than the PLAGA microspheres; however, the stainlesssteel mold, used more often for the composite microspheres, dissipatescharge buildup more readily than the PTFE mold, which is used more oftenfor the PLAGA microspheres, possibly explaining why there is asignificant loss for Phase B (p<0.05). Mesh for Phase A is notsusceptible to this loss.

Compressive modulus and yield strength were obtained for seeded andacellular control scaffolds at days 0, 7, 21, and 35 of culture. A rapiddecrease in compressive modulus was observed following day 0, possiblydue to rapid initial polymer degradation. By day 35, the seededscaffolds exhibited a greater compressive modulus (FIG. 51A) and yieldstrength (FIG. 51B), possibly due to cellular extracellular matrix andmineralization compensating loss of scaffold strength due to polymerdegradation.

In this experiment, the cell types were switched from bovine ACLfibroblasts and trabecular bone osteoblast-like cells to human hamstringtendon fibroblasts and trabecular bone osteoblasts due to the increasedclinical relevance of these new cell types. This experiment aimed toacquire quantitative data about cell proliferation and migrationthroughout the three phases, as well as cellular alkaline phosphataseactivity in each phase of the scaffold.

Based on the previous experiment performed with bovine cells, it isapparent that the biomimetic, multi-phased scaffolds support the growthand ECM production of both osteoblasts and fibroblasts. After 28 days ofculture, collagen production was evident in all three phases andmineralized matrix was found in the bone and interface regions.Osteoblast and fibroblast interaction at the interface (Phase B)suggests that these cells may play a significant role in the developmentof a functional insertion site. These findings demonstrate that thisnovel scaffold is capable of simultaneously supporting the growth ofmultiple cell types and can be used as a model system to regenerate thesoft tissue to bone interface. Additional studies can focus on scaffoldoptimization and the development of the interface on the novel scaffold.

Example 10 Electrospun PLAGA Meshes in the Multi-Phased Scaffold Design

The objective of the set of experiments was to incorporate electrospunPLAGA meshes into the multi-phased scaffold design, substituting theEthicon mesh phase, and allowing the entire scaffold to be madein-house.

Electrospinning, short for electrostatic spinning, is a relatively newterm that describes a principle first discovered in the first half ofthe 20th century (see, for example, U.S. Pat. Nos. 1,975,504, 2,160,962,2,187,306, 2,323,025 and 2,349,950 to Formhals, the entire contents ofwhich are incorporated herein by reference). Electrostatic spinninginvolves the fabrication of fibers by applying a high electric potentialto a polymer solution. The material to be electrospun, or dissolved intoa solution in the case of polymers, is loaded into a syringe or spoon,and a high potential is applied between the solution and a groundedsubstrate. As the potential is increased, the electrostatic forceapplied to the polymer solution overcomes surface tension, distortingthe solution droplet into a Taylor cone from which a jet of solution isejected toward the grounded plate. The jet splays into randomly orientedfibers, assuming that the solution has a high cohesive strength, linkedto polymer chain molecular weight, to prevent droplets from forminginstead of fibers in a process known as electrospraying. These fibershave diameters ranging from nanometer scale to greater than 1 μm and aredeposited onto the grounded substrate or onto objects inserted into theelectric field forming a non-woven mesh. Mesh characteristics can becustomized by altering electrospinning parameters. For example, fiberdiameter and morphology can be altered, including the formation of beadsalong the fibers, by controlling applied voltage and polymer solutionsurface tension and viscosity. Also, fiber orientation can be controlledby rotating the grounded substrate. This high degree of customizabilityand ability to use many different materials, such as biodegradablepolymers and silks, grant this fabrication method a high potential inthe development of materials for biomedical application. Management offiber diameter allows surface area to be controlled, and polymers withdifferent degradation rates can be combined in various ratios to controlfiber degradation, both of which are significant in drug deliveryapplications. Also, controlling the orientation of fiber depositiongrants a degree of control over cell attachment and migration. Moreover,the ability to electrospin fiber meshes onto non-metal objects placed inthe electric field enables the fabrication of multiphasic scaffoldsystems.

Here, in order to obtain precise parameters for the mesh fibers,including fiber diameter, morphology, and alignment, the effects ofprocessing parameters on fiber characteristics were studied. Avariable-speed rotating drum was designed and constructed to serve as asubstrate for aligned fibers, and Theological experiments were performedon the polymer solutions to determine the effect of polymerconcentration on solution viscosity and the subsequent effect ofsolution viscosity on fiber diameter and morphology.

In addition to determining the speed of each gear, the effect of eachspeed on fiber alignment was determined qualitatively. A 30% v/v PLAGAsolution was prepared with 60% dimethylformamide and 10% ethanol, andthis solution was electrospun onto the rotating drum at each of the fourspeed settings. The resulting meshes were examined by scanning electronmicroscopy (JEOL 5600LV).

The relationship between polymer concentration (Alkermes 85:15 PLAGA)and solution viscosity was determine by means of a Theological study.Three concentrations of polymer were tested-20%, 30%, and 40% v/v-indimethylformamide (DMF) and ethanol. The composition of each solution islisted in Table 12 shown below.

TABLE 12 PLGA DMF EtOH 1 20% 70% 10% 2 30% 60% 10% 3 40% 50% 10%

Solutions were analyzed using an Advanced Rheometer AR 2000t. There wasvariability in the viscosity measurements (n=1) at different strainrates due to the evaporation of solvent during testing. The geometryused for the viscosity measurements was a 25 mm stainless steel disc. Asolvent trap was not used since it is not designed to fit with thisgeometry and a prior trial using the solvent trap with another geometryresulted in poor results, possibly because water from the solvent trapseal interacted with the polymer solution. Additional trials can use asolvent trap to obtain consistent and reliable values for viscosity. Forthe present study, averages were taken of the viscosity measurementstaken at strain rates tested after the equipment had equilibrated. As aresult, there are standard deviations for the viscosity measurementseven with an n of 1.

The surface velocity of the rotating drum was seen to increase withincreased pulley positions from gear 1 to gear 4, as shown in Table 13below. The degree of fiber alignment increased with increasing drumvelocity, as seen in the SEMs of each mesh (see FIGS. 52A-D).

TABLE 13 Gear RPM Vs (m/s) 1^(st) 23.3 7.4 2^(nd) 29.5 9.4 3^(rd) 46.215 4^(th) 63.4 20

It was found that (as expected) the degree of fiber orientationincreased with increasing drum rotational velocity. The image wasanalyzed and a histogram of fiber angles was generated against thehorizontal axis of the image at regular interval across the image. Thus,the degree of alignment of the fibers can be quantified. It is desirableto control the degree of fiber alignment in the electrospun meshes sothat the extracellular environment found at the interface can bemimicked. By producing biomimetic scaffolds, it was intended to directcell growth to reproduce the tissue inhomogeneity found at the nativeACL insertions. In addition to controlling the fiber alignment, it isdesirable to control fiber diameter and morphology. It was previouslydetermined that substituting 10% of the DMF in the polymer solutionswith ethanol reduces the surface tension of the solution and results ina significant reduction in the number of beads formed along the fiberswhen electrospinning PLAGA. This effect was also observed by Fong etal., who reduced the number of beads in electrospun polyethylene oxide)(PEO) meshes by the addition of ethanol. Surface tension of the polymersolution acts to form spheres during the electrospinning process. Byreducing the solution surface tension, the formation of spheres is lessfavorable and straighter fibers result. Fong et al. also determined thatthe addition of ethanol increased the viscosity of the PEO:watersolutions, which also favors the formation of straight fibers, andresults in increased fiber diameter. Deitzel et al. also havedemonstrated a relationship between PEO:water solution viscosity andfiber diameter, with fiber diameter increasing with increasing viscosityaccording to a power law. A relationship between solution viscosity andconcentration of polymer can be determined in order to understand howPLAGA:N, N-DMF viscosity affects fiber diameter and morphology. Theeffect of solution viscosity on fiber diameter and morphology can bedetermined by spinning the various solutions and examining the resultingmeshes by SEM. Other variables can affect the fiber parameters. Bychanging the percentage of polymer, the surface tensions of the polymersolutions also change in addition to the viscosity. Therefore, inaddition to testing the viscosities of each solution, the surfacetension of each solution are measured. It is desirable to keep allvariables constant except for viscosity in order to truly determine theeffect of solution viscosity on fiber characteristics. However, theinterrelation of many of the electrospinning parameters complicates theprocess.

A PLAGA mesh was electrospun directly onto a microsphere scaffold.

This is one way to incorporate the mesh. In addition, the scaffolds canbe secured to the drum and aligned fibers electrospin directly onto thescaffolds.

However, because of the high rotational velocities, it is difficult tosecure the scaffolds and prevent them from flying off the drum when itbegins rotating. Alternatively, aligned fiber meshes can simply be spunseparately, and then later sintered to the microsphere scaffolds. Forexample, aligned fiber meshes can be electrospun onto aluminum foil,then wrapped around a rod with multiple mesh sheets sintered together toobtain a hollow cylinder of aligned fibers.

FIGS. 52E and 52F show scanning electron microscopy (SEM) images ofanother embodiment of multi-phased scaffold, with 85:15 PLAGAelectrospun mesh joined with PLAGA:BG composite microspheres.

Example 11 A Fully Synthetic Implantable Scaffold

With reference to FIG. 55, the fully synthetic implantable scaffoldaccording to the present invention may be made according to thefollowing procedure. A synthetic graft (Phase A) is made usingconventional techniques from any suitable bioresorbable and/orbiodegradable synthetic polymer material, such as for example, aliphaticpolyesters, poly(amino acids), copoly(ether-esters), polyalkylenesoxalates, polyamides, poly(iminocarbonates), polyorthoesters,polyoxaesters, polyamidoesters, poly(ε-caprolactone)s, polyanhydrides,polyarylates, polyphosphazenes, polyhydroxyalkanoates, polysaccharides,degradable polyurethanes, and biopolymers, and a blend of two or more ofthe preceding polymers. For example, the synthetic graft may be madefrom at least one of a poly(lactide-co-glycolide), poly(lactide) andpoly(glycolide).

Phases B and C of the scaffold are made from any suitable material aspreviously defined herein. For example, phase C may be constructed of apolymer-ceramic composite with high Ca—P and phase B may be constructedof a polymer only or a polymer-ceramic composite with a lower Ca—P thanphase C.

The fully synthetic scaffold is assembled by first attaching Phase B toboth ends of Phase A using conventional techniques, such as sintering orsolvent evaporation. The Phase C is attached to the free end of eachrespective Phase B, again using conventional techniques, such assintering or solvent evaporation. Thus formed, the scaffold is amulti-phased, biodegradable, and osteointegrative composite having thestructure: Phase C-Phase B-Phase A-Phase B-Phase C.

As previously described, each phase of the scaffold may be furthertreated to initiate, promote, and enhance integration of the scaffoldwith anatomical environment to which it is delivered. Such treatment mayinclude seeding appropriate phases of the scaffold with appropriate celltypes (e.g., seeding a phase C which will be inserted into bone withosteoblasts, osteoblast-like cells, and/or stem cells or seeding phaseC, which will be attached to soft tissue with fibroblasts, chondrocytesand/or stem cells), treating appropriate phases of the scaffold with amedicament, such as for example anti-infectives, antibiotics,bisphosphonate, hormones, analgesics, anti-inflammatory agents, growthfactors, angiogenic factors, chemotherapeutic agents, anti-rejectionagents, and RGD peptides.

The fully synthetic implantable scaffold according to the presentinvention may be used in any ligament or tendon repair orreconstruction. Preferably, the fully synthetic implantable scaffoldaccording to the present invention is used to repair or reconstructanterior cruciate ligaments.

Example 12 The Triphasic Scaffold for Use as Graft Collar orInterference Screw

Clinically, the hamstring tendon graft is mechanically fixedextra-articularly by looping the graft around a transfemoral pin in thefemoral bone tunnel, while a screw with a washer or a staple is used tofix the graft to the tibia. Interference screws have been used in thebone tunnel, but with limited success due to graft laceration and poorfixation strength. With mechanical fixation, the fibrocartilageinterface is not regenerated after ACL reconstruction. Anon-physiologic, fibrovascular scar tissue is instead formed within thebone tunnel as part of the healing process. The presence of thispartially mineralized layer within the tunnel renders the graft-bonefixation site the weakest point mechanically (Rodeo, 2006). This problemis exacerbated by the active lifestyle of ACL injury patients (15-35years old), which necessitates higher fixation strength and expeditedhealing. Thus, graft-to-bone fixation remains a significant clinicalproblem.

The subject approach to addressing the challenge of biological fixationis original and represents a significant departure from the conventionalfocus on tendon-to-bone healing within the bone tunnel. It is emphasizedhere that the native anatomical fibrocartilage interface is orthogonalto the subchondral bone and continuous with surrounding articularcartilage. In addition, the neo-fibrocartilage formed within the bonetunnel represents the mechanical weak link for tendon-to-boneintegration. Biological fixation therefore requires that the anatomicalfibrocartilage insertion is regenerated between graft and bone,accompanied by the complete mineralization of the tendon within the bonetunnel.

It is envisioned that the triphasic scaffold may be used clinically aseither as a graft collar or an interference screw during ACLreconstruction surgery. The ultimate goal is to facilitate the formationof the anatomic fibrocartilage interface directly on the soft tissuegraft. As a graft collar, the scaffold will be fabricated as a hollowcylinder through which the ACL graft can be inserted. As shown in FIG.62, the collar can be sutured or secured to the ends of the tendongraft. Fixation is achieved by inserting the collar-graft complex intothe bone tunnel, with Phase C positioned inside the bone tunnel, Phase Bflush with articular cartilage, and only Phase A directly exposed to thejoint cavity. It is anticipated that the designed heterogeneity andoptimized interaction between MSC-derived cells will induce theformation of a fibrocartilage interface directly onto the graft. Graftintegration within the bone tunnel will be facilitated by Phase C, theosteointegrative polymer-ceramic composite, and with the eventualaddition of growth factors (e.g., bone morphogenetic proteins), whichwill induce osteointegration and mineralization of the tendon graftwithin the bone tunnel.

For use as an interference screw, the triphasic scaffold can befabricated as matching portions of the hollow cylinder, with eachportion containing the three scaffold phases. As shown in FIG. 63, thetwo matching portions will encase the soft tissue graft on all sides.The relative position of each phase of the triphasic scaffold would bein the anatomical position, i.e., with Phase A (soft tissue) exposed tothe joint cavity, Phase B (fibrocartilage interface) flush witharticular cartilage, and Phase C (bone) encased within the bone tunnel.There are several advantages to this novel interference screw design: 1)the biomimetic triphasic screw design enables the regeneration of therelevant tissue types on the scaffold system, 2) the partitioned designpermits the application of mechanical loading to the graft, which hasbeen known to induce fibrocartilage formation, and 3) the tendon graftis in contact with the triphasic scaffold on all sides. Any appliedmechanical and chemical stimulation would be uniformly experienced bythe graft.

The optimal outcome scenario post-degradation of the screw or graftcollar is to have a completely mineralized tissue within the bonetunnel, accompanied by the formation of a physiologically equivalentfibrocartilage insertion directly outside the bone.

For ligament tissue engineering, the triphasic scaffold may be coupledwith synthetic grafts for ACL replacement. The future design of ACLreplacement grafts must take into consideration the integration of thegraft with bone. In this integrative ACL prosthesis design, the ACLprosthesis will contain three regions, a bony end consisting of Phase C,followed by Phase B, then by polymer fiber-based ACL portion. Thetriphasic scaffold can also be incorporated into any existing ACLprosthesis design, as the soft tissue graft shown in FIGS. 62 and 63 caneasily be replaced by any synthetic ACL reconstruction scaffold. Forexample, in the case of a degradable polymer-based ACL prosthesis(Cooper, 2005), the triphasic scaffold can be sintered onto the polymerscaffold and implanted for ACL reconstruction.

One common feature in the above examples of clinical application is thefocus on engineering soft tissue-to-bone integration ex vivo, whichwould reduce the complexity of graft reconstruction to just bone-to-boneintegration in vivo. This is more feasible clinically as it is much moredifficult to integrate soft tissue with bone compared to bone-to-boneintegration.

Extensive characterization of the chemical and mechanical properties ofthe interface (Wang, 2006; Spalazzi, 2006; Moffat, 2005) has beenconducted and novel in vitro co-culture (Wang, 2005) and tri-culture(Wang, 2006) models have been developed to examine the role of cell-cellinteractions in interface formation. In combination with knowledge of invivo models of tendon-bone healing (Rodeo, 1993; Kawamura, 2005), thereis a solid foundation and clear rationale for the described approach.

From a broader impact perspective, this approach is also unique in thatprevious tissue engineering methods have focused predominantly on thedesign of a single type of tissue (e.g., only ligament or bone) on ascaffold with uniform properties, when the application may have involvedmore than one tissue type. Moreover, the novel scaffold design andco-culture methods described here can be applied to treat other clinicalconditions (e.g., rotator cuff, osteoarthritis) and will enable thedesign of a new generation of integrative fixation devices. Thedescribed studies will also provide fundamental insights into themechanism of soft tissue-bone interface regeneration.

Clinical feasibility of the scaffold was determined by testing thehypothesis that the biomimetic matrix heterogeneity engineered on thetriphasic scaffold will be maintained in vivo in an intra-articularmodel. A summary schematic of this research approach is presented belowin FIG. 64. It was determined that modifications to the scaffold designwere necessary to achieve distinct cell and matrix regions in vivo.

Scaffold Design Optimization

Based on the outcomes of in vitro and in vivo co-culture and tri-cultureexperiments, the multi-phased scaffold design has been improved upon,with the goal of localizing the interface-relevant cells within Phase Bwithout compromising the scaffold design requirements (higher porosityand pore diameter) necessary for Phase A. Specifically, a degradablecell barrier between adjacent phases has been incorporated. This barrieris based on a polylactide-co-glycolide (PLGA) electrospun nanofiber mesh(FIG. 65-I), which, based on porosimetry analysis, has an average porediameter of 5.2±0.9 μm. This nanofiber mesh will prevent unwanted cellmigration and gel infiltration into Phase A or Phase C. Celllocalization is important as 3-D co-culture results demonstrate thatcell-specific distribution is required for the development of thebiomimetic, controlled matrix distribution on the multi-phased scaffold.

Preliminary cell tracking results of fibroblasts and osteoblaststri-cultured with chondrocytes loaded in hydrogel for 24 hours on themodified scaffold are shown in FIG. 65-II. Fibroblasts, chondrocytes,and osteoblasts were detected only in their respective phases asdetermined by fluorescence confocal microscopy. The nanofiber meshserved as an effective barrier to gel infiltration and unwanted cellcross-migration. It is anticipated that the mesh will degrade over time,having ensured the establishment of cell-specific regions intri-culture.

Mechano-Actuve Scaffold Induces Remodeling and Expression ofFibrocartilage Markers on Tendon Grafts

Based on the hypothesis that mechanical loading, in addition tomultiphasic scaffold design and heterotypic cellular interactions, willbe required for interface generation, this experiment focuses on thedesign and evaluation of a novel mechano-active scaffold that is capableof applying compression to tendon grafts and inducing metaplasia oftendon into fibrocartilage. Specifically the novel scaffold systemcombines a degradable graft collar with nanofiber meshes fabricated frompoly(latic-co-glycolic acid) (PLGA). One objective of the experiment isto characterize the contractile properties of the nanofiber mesh as wellas the mesh and scaffold complex. A second objective of the experimentis to evaluate the effect of scaffold induced compression onfibrocartilage development on tendon graft, focusing on matrixremodeling and the development of fibrocartilage-related markers.

First, aligned nanofiber meshes were fabricated by electrospinning. Aviscous polymer solution consisting of 35% poly(D,L-lactic-co-glycolicacid) 85:15 (PLGA, I.V.=0.70 dL/g, Lakeshore Biomaterials, Birmingham,Ala.), 55% N,N-dimethylformamide (Sigma, St. Louis, Mo.), and 10%ethanol (Commercial Alcohol, Inc., Toronto, Ontario) was loaded into asyringe fitted with an 18-gauge needle (Becton Dickinson, FranklinLakes, N.J.). Aligned fibers were obtained using an aluminum drum withan outer diameter of 10.2 cm rotating with a surface velocity of 20 m/2.A constant flow rate of 1 mL/hr was maintained using a syringe pump(Harvard Apparatus, Holliston, Mass.), and an electrical potential wasapplied between the needle and the grounded substrate (distance=10 cm)using a high voltage DC power supply (Spellman, Hauppauge, N.Y., 8-10kV). Fiber morphology, diameter and alignment of the as-fabricated meshsamples were analyzed using scanning electron microscopy (SEM). Thesamples were sputter-coated with gold (LVC-76, Plasma Sciences, Lorton,Va.) and subsequently imaged (JSM 5600LV, JEOL, Tokyo, Japan) at anaccelerating voltage of gkV.

The nanofiber mesh exhibited a high degree of alignment with an averagefiber diameter of 0.9±0.4 μm. Anisotropic mesh contractile behavior wasobserved in the mesh, with significantly higher contraction found in thedirection of nanofiber alignment. Specifically, the mesh contracted over57% along the aligned fiber direction (y-axis) by 2 hours, with lessthan 13% reduction in the x-axis. Mesh contraction continued over time,exhibiting over 70% contraction in the y-axis and 20% in the x-axis by24 hours and stabilizing thereafter, with no significant differencesfound between the 24-and 72-hour groups.

The said nanofiber is then wrapped around a tendon graft collar based ona sintered microsphere scaffold fabricated following published methods.In addition, patellar tendon grafts were isolated from neonatal bovinetibiofemoral joints obtained from a local abattoir. The compression ofthe nanofiber mesh, the graft collar scaffold with nanofiber mesh, thetendon with nanofiber mesh, and finally the tendon graft with the graftcollar scaffold and the nanofiber mesh were evaluated. Further, theeffects of compression on graft cellularity, organization, matrixcontent, and cell phenotype were evaluated.

It was found that complex of nanofiber mesh and graft collar was able toapply a physiological range of compressive loading on the tendons.Moreover scaffold-mediated compression promoted matrix remodeling,maintained graft glycosaminoglycan content and induced gene expressionfor fibrocartilage markers, including type II collagen, aggrecan coreprotein, and TGF-β3.

Mesenchymal Stem Cells and Differentiation into Interface-Relevant CellPopulations

The experiments will also utilize fibroblasts, chondrocytes, andosteoblasts derived from adult mesenchymal stem cells (MSCs) originatedfrom human bone marrow. The MSCs are chosen because they are ideal fortissue engineering applications. These cells can be harvested from thepatient prior to surgery, expanded, and pre-differentiated into desiredcell types, and then seeded onto 3-D scaffolds. In addition to beingautologous, MSCs can differentiate into fibroblasts (Pittenger, 1999;Moreau, 2005), chondrocytes (Pittenger, 1999; Meinel, 2004), andosteoblasts (Pittenger, 1999; Mauney, 2005) which are the relevant celltypes found at the soft tissue-bone interface. This versatility willsimplify the tissue harvest process to a single procedure instead of thenormal three required to obtain the three types of cells. Successfulimplementation of MSC-derived cells will significantly enhance theclinical feasibility and translational potential of the triphasicscaffold.

Specifically, MSCs purchased from Cambrex will be pre-differentiatedinto fibroblasts (Fb), chondrocytes (Ch), and osteoblasts (Ob) based onwell-established protocols. The fibrogenic media will contain 1 ng/mL ofbasic fibroblast growth factor, 5 ng/mL of transforming growthfactor-beta (TGF-β3) and 50 μg/ml of L-Ascorbic Acid-2-Phosphate (AA)(Moreau, 2005; Altman, 2002). The chondrogenic media will contain 5ng/mL TGF-β3, 0.1 mM non-essential amino acids, 50 μg/ml AA, 10 nMdexamethasone (Dex), and 5 μg/ml of insulin16]. The osteogenic mediawill contain 10 nM Dex, 10 mM of β-glycerophosphate, and 50 μg/ml AA(Mauney, 2005).

Intra-Articular ACL Reconstruction Model

The study will use male athymic rats (Charles River Laboratories, meanweight 300 grams) to demonstrate unilateral ACL reconstruction (Rodeo,2006) using a flexor digitorum longus tendon graft from the ipsilaterallimb, as shown in FIG. 66-I. The rats will be anesthetized with amixture of ketamine hydrochloride 80 mg/kg and xylazine 5 mg/kg,administered intraperitoneally. Ampicillin 25 mg/kg subcutaneousinjection will be used for antibiotic prophylaxis. After appropriateanesthesia, the rat will be prepared for sterile surgery. The flexordigitorum longus tendon will be harvested via a longitudinal incisionmade on the medial aspect of the distal leg and ankle. The full lengthof the flexor digitorum longus tendon (average length 20 mm) will beharvested. An incision will be made over the rat knee, and a lateralparapatellar arthrotomy will be performed. The ACL, PCL, MCL, and LCLwill be excised. Sectioning these ligaments causes minimal trauma to theknee and is not expected to affect the overall biologic response thatwill already occur from the knee arthrotomy. Using a needle with outerdiameter of 2.5 mm, a bone tunnel will be made in the proximal tibia andthe distal femur, entering the joint at the attachment sites of the ACL.We will measure the total length of the femur-tendon-tibia complex todetermine the amount of displacement required to apply 1% and 10%strain.

The triphasic scaffold fabricated in the form of the graft collar willbe used for implantation. After incorporating the graft collar onto theflexor tendon graft, the graft-scaffold complex will be passed throughthe bone tunnels to replace the ACL. Both ends of the grafted tendonwill be secured to the surrounding periosteum at the extra-articulartunnel exit sites at the distal femur and well as proximal tibia using4-0 Ethibond suture. Post-operative activity will be controlled using anexternal fixator that we have designed and fabricated for rat knees(Rodeo, 2006).

Cell Tracking In Vivo

A further objective of these experiments is to track the three types ofimplanted cell populations in vivo and to determine their presence overa 4-week implantation period. Cell Labeling—After pre-differentiation ofMSCs into Fb, Ch, and Ob, cells will be seeded based on the optimal cellseeding density (cells/cm³) on their designated phase of the triphasicscaffold based on results from Phase I. As shown in FIG. 65, the Fb willbe pre-labeled with Vybrant DiD dye (green), Ch with Vybrant DiO (red),and Ob with Vybrant Dil (yellow). All dyes can be purchased fromMolecular Probes. The pre-label cells will be seeded on their respectivephases of the triphasic scaffold collar, and tricultured for 2 daysfollowing established protocols (Spalazzi, 2006). As summarized in FIG.67, the scaffold (n=3 per group) will be implanted for 1, 2, and 4weeks, and the presence of the cells will be tracked over time andcorrelated to the formation of fibrocartilage tissue on the triphasicscaffold. At each time point, the scaffold collar+graft complex will beexcised and cryosectioned for fluorescence microscopy (cell imaging) andhistological analysis (fibrocartilage formation). Specifically,development of interface-relevant markers will be determined:proteoglycan and mineral deposition, as well as immunohistochemistry forcollagen types I, II, III, IX, and X. Acellular scaffolds and unoperatedcontralateral insertion sites will serve as additional controls. A totalof 45 animals (15 per time point) will be needed for this experiment.

In Vivo Evaluation for Interface Regeneration

This experiment further focuses on interface regeneration on thetri-cultured, triphasic scaffold in an intra-articular ACLreconstruction model. Specifically, MSC-derived fibroblasts,chondrocytes and osteoblasts will be seeded on their respective phasesof the triphasic scaffold, and cultured in vitro for 2 days (Spalazzi,2006). The scaffold will be implanted following the experimental designoutlined in FIG. 68. Each animal will receive one scaffold (randomlyselected) and will be sacrificed at 4, 8, and weeks. Outcomes will beevaluated using histomorphometric, micro-CT, and biomechanical analyses.Quantitative histomorphometric measurements will be made using theBioquant Image Analysis system (R&M Biometrics, Inc., Nashville, Tenn.)available in the Analytical Microscopy Laboratory (Director, Dr. S.Doty). The implant evaluation methods successfully utilized in thepreviously described in vivo studies will also be used here.Specifically, the development of a fibrocartilage-like tissue andinterfacial markers (n=3) will be determine. Scaffold mechanicalproperties (n=6) will also be determined over time. Mineralization(total bone mineral content, bone volume fraction, and mineraldistribution) will be analyzed by micro-CT prior to mechanical testing,so an additional sample is not needed. A push-out test (Knowles, 1992)will be performed on week 12 samples (tri-culture only, n=6) in order todetermine the osteointegration potential of Phase C within the bonetunnel. A total of 168 male athymic nude rats (54 animals each for weeks4 & 8, and 60 animals for week 12) will be used in this experiment.

Expected Outcomes

It is anticipated that for the in vivo cell tracking experiment, allthree cell types will persist at the implantation site for up to 4weeks, and that the seeded chondrocytes will contribute to the formationof a fibrocartilage-like region on the interface phase (Phase B) of thetriphasic scaffold. For the in vivo evaluation of interface regenerationexperiment, it is expected that an interface-like region will form onthe scaffold post-ACL reconstruction.

In these experiments, the formation of a fibrocartilage-like tissue onthe interface phase of the triphasic scaffold has been focused on forseveral reasons. The long term role of the scaffold as a graft collar isto induce fibrocartilage formation on the reconstructed graft. Afterestablishing the stability of the triphasic scaffold in theintra-articular model, and the viability of application of controlledmechanical stimulation to induce fibrocartilage formation on the graft,the next stage of the project will focus on the application ofcontrolled chemical stimulation to induce fibrocartilage formation onthe graft. For example, phase-specific growth factor delivery can beincorporated to provide chemical stimuli for interface regeneration. Itis however critical to first establish the feasibility of thetri-culture, triphasic scaffold in a physiologically relevantintra-articular model.

REFERENCES

The following documents, as well as those cited within thisspecification, are specifically incorporated by reference to the extentthat they provide or teach exemplary methodology, techniques and/orcompositions supplemental to those employed herein.

-   1. Abate, J A, Fadale, P D, Hulstyn, M J & Walsh, W R, (1998)    “Initial fixation strength of polylactic acid interference screws in    anterior cruciate ligament reconstruction,” Arthroscopy 14:278-284.-   2. Albro, M B, Chahine, N O, Caligaris, M, Wei, V I,    Likhitpanichkul, M, Ng, K W, Hung, C T & Ateshian, G A, (2007)    “Osmotic loading of spherical gels: a biomimetic study of hindered    transport in the cell protoplasm,” J. Biomech. Eng. 129:503.-   3. Allum, R L, (2001) “BASK Instructional Lecture 1: graft selection    in anterior cruciate ligament reconstruction,” Knee 8:69-72.-   4. Altman, G H, et al. (2002) “Advanced bioreactor with controlled    application of multidimensional strain for tissue engineering,”    Journal of Biomechanical Engineering 124:742-749.-   5. Altman, G H, et al., (2002) “Silk matrix for tissue engineered    anterior cruciate ligaments,”, 23:4131-4141.-   6. American Academy of Orthopaedic Surgeons, (1997) “Arthoplasty and    Total Joint Replacement Procedures: United States 1990 to 1997,”    (Report, United States).-   7. Anderson, K, et al., (2001) “Augmentation of tendon healing in an    intraarticular bone tunnel with use of a bone growth factor,” Am. J    Sports Med. 29:689-698.-   8. Badami, A S, Kreke, M R, Thompson, M S, Riffle, J S & Goldstein,    A S, (2006) “Effect of fiber diameter on spreading, proliferation,    and differentiation of osteoblastic cells on electrospun poly(lactic    acid) substrates,” Biomaterials, 27:596.-   9. Badylak, S F (2002) “The extracellular matrix as a nanofiber    scaffold for tissue reconstruction,” Semin. Cell Dev. Biol. 13:377.-   10. Baker, B M and Mauck, R L, (2007) “The effect of nanofiber    alignment on the maturation of engineered meniscus constructs,”    Biomaterials, 28:1967.-   11. Bashur, C A, Dahlgren, L A & Goldstein, A S (2006) “Effect of    fiber diameter and orientation on fibroblast morphology and    proliferation on electrospun poly(D,L-lactic-co-glycolic acid)    meshes,” Biomaterials, 27:5681.-   12. Batycky, R P, Hanes, J, Langer, R & Edwards, D A, (1997) “A 1.5    theoretical model of erosion and macromolecular drug release from    biodegrading microspheres,” J. Pharmaceutical Sciences 86:1464-1477.-   13. Bellincampi, L D, Closkey, R F, Prasad, R, Zawadsky, J P & Dunn,    M G, (1998) “Viability of fibroblast-seeded ligament analogs after    autogenous implantation,” J. Orthop. Res. 16:414-420.-   14. Benjamin, M, Evans, E J, Copp, L, (1986) “The histology of    tendon attachments to bone in man,” J Anat. 149:89-100.-   15. Benjamin, M, Evans, E J, Rao, R D, Findlay, J A & Pemberton, D    J, (1991) “Quantitative differences in the histology of the    attachment zones of the meniscal horns in the knee joint of man,” J.    Anat. 177:127-134.-   16. Benjamin, M, Kumai, T, MiIz, S, Boszczyk, B M, Boszczyk, A A,    Ralphs, J R, (2002) “The skeletal attachment of tendon-tendon    “entheses”,” Comp Biochem. Physiol A. MoI. Integr. Physiol.    133(4):931-945.-   17. Benjamin, M, Ralphs, J R (1998) “Fibrocartilage in tendons and    ligaments—an adaptation to compressive load,” J. Anat. 193(Pt    4):481-494.-   18. Berg, E E, (1996) “Autograft bone-patella tendon-bone plug    comminution with loss of ligament fixation and stability,”    Arthroscopy 12:232-235.-   19. Beynnon, B D, et al., (1996) “A sagittal plane model of the knee    and cruciate ligaments with application of a sensitivity    analysis,” J. Biomech. Eng. 118:227-239.-   20. Beynnon, B D, et al., (1997) “The effect of functional knee    bracing on the anterior cruciate ligament in the weightbearing and    nonweightbearing knee,” Am. J. Sports Med. 25:353-359.-   21. Beynnon, B D, et al., (2002) “Anterior Cruciate Ligament    Replacement: Comparison of Bone-Patellar Tendon-Bone Grafts with    Two-Strand Hamstring Grafts,” J Bone Joint Surg Am, 84-A:1503-1513.-   22. Blevins, F T, Djurasovic, M, Flatow, E L, Vogel, K G (1997)    “Biology of the rotator cuff tendon,” Orthop. Clin. North Am.    28(1):1-16.-   23. Blickenstaff, K R. Grana, W A & Egle, D, (1997) “Analysis of a    semitendinosus autograft in a rabbit model,” Am. J. Sports Med.    25:554-559.-   24. Bolton, C W & Bruchman, W C, (1985) “The GORE-TEX expanded    polytetrafluoroethylene prosthetic ligament. An in vitro and in vivo    evaluation,” Clin. Orthop. 202-213.-   25. Bonfield, W, (1988) “Composites for bone replacement,” J.    Biomed. Eng. 10:522-526.-   26. Borden, M, Attawia, M, Khan, Y & Laurencin, C T, (2002) “Tissue    engineered microsphere-based matrices for bone repair: design and    evaluation,” Biomaterials, 23:551-559.-   27. Boskey, A L, et al., (1996) “The mechanism of    beta-glycerophosphate action in mineralizing chick limb-bud    mesenchymal cell cultures,” J. Bone Min. Res. 11:1694-1702.-   28. Brady, G A, Eisinger, M, Arnoczky, S P & Warren, R F, (1988) “In    vitro fibroblast seeding of prosthetic anterior cruciate ligaments.    A preliminary study,” Am. J. Sports Med. 16:203-208.-   29. Brand, J, Jr., Weiler, A, Caborn, D N, Brown, C H, Jr. &    Johnson, D L (2000) “Graft fixation in cruciate ligament    reconstruction,” Am. J. Sports Med. 28:761-774.-   30. Brody, G A, Eisinger, M, Arnoczky, S P & Warren, R F, (1988) “In    vitro fibroblast seeding of prosthetic anterior cruciate ligaments.    A preliminary study,” Am. J. Sports Med. 16:203-208.-   31. Bromage, T G, Smolyar, I, Doty, S B, Holton, E & Zuyev, A    N, (1998) “Bone growth rate and relative mineralization density    during space flight,” Scanning 20:238-239.-   32. Burkart, A, Imhoff, A B & Roscher, E, (2000) “Foreign-body    reaction to the bioabsorbable suretac device,” Arthroscopy,    16:91-95.-   33. Butler, D L, Goldstein, S A & Guilak, F, (2000) “Functional    tissue engineering: the role of biomechanics,” J. Biomech. Eng.    122:570-575.-   34. Chen, C H, et al., (2003) “Enveloping the tendon graft with    periosteum to enhance tendon-bone healing in a bone tunnel: A    biomechanical and histologic study in rabbits,” Arthroscopy    19:290-296.-   35. Cheung, H S & McCarty, D J, (1985) “Mitogenesis induced by    calcium-containing crystals. Role of intracellular dissolution,”    Exp. Cell Res. 157:63-70.-   36. Christenson, E M, et al. (2007) “Nanobiomaterial applications in    orthopedics,” J. Orthop. Res. 25:11-22.-   37. Chun I, et al. (1995) “Fine Fibres Spun By Electrospinning    Process From Polymer Solution And Polymer Melts.” Dissertation, The    University of Akron.-   38. Clark, J M & Sidles, J A, (1990) “The interrelation of fiber    bundles in the anterior cruciate ligament,” J. Orthop. Res.    8:180-188.-   39. Codman, E, (1934) “The Shoulder, Rupture of the Supraspinatus    Tendon and Other Lesions In or About the Subacromial Bursa,” Thomas    Todd, BostoN.-   40. Cole, B J, EIAttrache, N S & Anbari, A, (2007) “Arthroscopic    rotator cuff repairs: an anatomical and biomechanical rationale for    different suture-anchor repair configurations,” Arthroscopy    23:662-669.-   41. Coons, D A and Alan, B F, (2006) “Tendon graft    substitutes-rotator cuff patches,” Sports Med. Arthrosc. 14:185-190.-   42. Cooper J A et al., (2005) “Fiber-based tissue-engineered    scaffold for ligament replacement: design considerations and in    vitro evaluation” Biomaterials, 26(13):1523-1532.-   43. Cooper, J A, (2002) “Design, optimization and in vivo evaluation    of a tissue-engineered anterior cruciate ligament replacement,”    Drexel University (Thesis/Dissertation) (2002).-   44. Cooper, J A, Lu, H H & Laurencin, C T, (2000) “Fiber-based    tissue engineering scaffold for ligament replacement: design    considerations and in vitro evaluation,” Proceedings of 5th World    Biomaterial Congress, 208 (Abstract).-   45. Costa, K D, Lee, E J & Holmes, J W, (2003) “Creating alignment    and anisotropy in engineered heart tissue: role of boundary    conditions in a model three-dimensional culture system,” Tissue Eng.    9:567-577.-   46. Courtney, T, Sacks, M S, Stankus, J, Guan, J & Wagner, W    R, (2006) “Design and analysis of tissue engineering scaffolds that    mimic soft tissue mechanical anisotropy,” Biomaterials,    27:3631-3638.-   47. Currey, J D, (1988) “The effect of porosity and mineral content    on the Young's modulus of elasticity of compact bone,” J Biomech,    21(2):131-139.-   48. Curtis, A and Wilkinson, C, (1997) “Topographical control of    cells,” Biomaterials, 18:1573-1583.-   49. Daniel, D M, et al., (1994) “Fate of the ACL-injured patient. A    prospective outcome study,” Am. J. Sports Med. 22:632-644.-   50. Deitzel et al., (2001) “Controlled deposition of electrospun    poly(ethylene oxide) fibers,” Polymer, 42(19):8163-8170.-   51. Deitzel et al., (2002) “Electrospinning of Nanofibers with    Specific Surface Chemistry,” Polymer, 43(3): 1025-1029.-   52. Dejardin, L M, Arnoczky, S P, Ewers, B J, Haut, R C & Clarke, R    B (2001) “Tissue engineered rotator cuff tendon using porcine small    intestine submucosa. Histologic and mechanical evaluation in dogs,”    Am. J. Sports Med. 29:175-184.-   53. Delon, I & Brown, N H (2007) “Integrins and the actin    cytoskeleton,” Curr. Opin. Cell Biol. 19:43-50.-   54. DeOrio, J K and Cofield, R H, (1984) “Results of a second    attempt at surgical repair of a failed initial rotator-cuff    repair,” J. Bone Joint Surg. Am. 66:563.-   55. Derwin, K A, Baker, A R, Spragg, R K, Leigh, D R & lannotti, J    P, (2006) “Commercial extracellular matrix scaffolds for rotator    cuff tendon repair. Biomechanical, biochemical, and cellular    properties,” J. Bone Joint Surg. Am. 88:2665.-   56. Drost, M R, Willems, P, Snijders, H, Huyghe, J M, Janssen, J D &    Huson, A, (1995) “Confined compression of canine annulus fibrosus    under chemical and mechanical loading,” J. Biomech. Eng. 11: 390.-   57. Ducheyne, P, (1987) “Bioceramics: material characteristics    versus in vivo behavior,” J. Biomed. Matls. Res. 21:219-236.-   58. Dunn, M G, Liesch, J B, Tiku, M L, Maxian, S H & Zawadsky, J    P, (1994) “The Tissue Engineering Approach to Ligament    Reconstruction,” Matls. Res. Soc. 331:13-18.-   59. El-Amin, S F, et al., (2001) “Human osteoblast integrin    expression on degradable polymeric materials for tissue engineered    bone,” J. Orthop. Res, 20(1):20-28.-   60. Engelmayr, G C, Jr., Papworth, G D, Watkins, S C, Mayer, J E,    Jr. & Sacks, M S, (2006) “Guidance of engineered tissue collagen    orientation by large-scale scaffold microstructures,” J. Biomech.    39:1819.-   61. Erben, R G, (1997) “Embedding of Bone Samples in    Methylmethacrylate: An Improved Method Suitable for Bone    Histomorphometry, Histochemistry, and Immunohistochemistry” J.    Histochem. Cytochem. 45:307-314.-   62. Ferguson, V L, Bushby, A J., Boyde, A, (2003) “Nanomechanical    properties and mineral concentration in articular calcified    cartilage and subchondral bone,” J. Anat. 203(2):191-202.-   63. Fleming, B, Beynnon, B, Howe, J, McLeod, W & Pope, M, (1992)    “Effect of tension and placement of a prosthetic anterior cruciate    ligament on the anteroposterior laxity of the knee,” J. Orthop. Res.    10:177-186.-   64. Fleming, B C, Abate, J A, Peura, G D & Beynnon, B D, (2001) “The    relationship between graft tensioning and the anterior-posterior    laxity in the anterior cruciate ligament reconstructed goat    knee,” J. Orthop. Res. 19:841-844.-   65. Fong et al., (1999) “Beaded nanofibers formed during    electrospinning” Polymer, 40:4585-4592.-   66. Formhals, A, (1934) “Process and Apparatus for Preparing    Artificial Threads,” U.S. Pat. No. 1,975,504, issued October, 1934.-   67. Fridrikh et al., (2003) “Controlling the fiber diameter during    electrospinning” Physical Review Letter, 90:144502.-   68. Fu, F H, Bennett, C H, Ma, C B, Menetrey, J & Lattermann,    C, (2000) “Current trends in anterior cruciate ligament    reconstruction. Part II. Operative procedures and clinical    correlations,” Am. J. Sports Med. 28:124-130.-   69. Fujihara, K, Kotaki, M & Ramakrishna, S (2005) “Guided bone    regeneration membrane made of polycaprolactone/calcium carbonate    composite nano-fibers,” Biomaterials, 26:4139.-   70. Fujikawa, K, Iseki, F & Seedhom, B B, (1989) “Arthroscopy after    anterior cruciate reconstruction with the Leeds-Keio ligament,” J.    Bone Joint Surg. Br. 71:566-570.-   71. Fung, Y C, (1972) “Stress-strain-history relations of soft    tissues in simple elongation. In Biomechanics: Its Foundations and    Objectives,” Fung, Y C, Perrone, N, Anliker, M, eds., Prentice-Hall:    San Diego, pp. 181-208.-   72. Galatz, L M, Ball, C M, Teefey, S A, Middleton, W D & Yamaguchi,    K, (2004) “The outcome and repair integrity of completely    arthroscopically repaired large and massive rotator cuff tears,” J.    Bone Joint Surg. Am. 86-A:219.-   73. Gao, J & Messner, K, (1996) “Quantitative comparison of soft    tissue-bone interface at chondral ligament insertions in the rabbit    knee joint,” J. Anat. 188:367-373.-   74. Gao, J, Rasanen, T, Persliden, J & Messner, K, (1996) “The    morphology of ligament insertions after failure at low strain    velocity: an evaluation of ligament entheses in the rabbit knee,” J.    Anat. 189:127-133.-   75. Garcia, A J (2005) “Get a grip: integrins in cell-biomaterial    interactions,” Biomaterials, 26:7525.-   76. Garreta, E, Gasset, D, Semino, C & Borros, S, (2007)    “Fabrication of a three-dimensional nanostructured biomaterial for    tissue engineering of bone,” Biomol. Eng. 24:75.-   77. Gartsman, G M (2001) “All arthroscopic rotator cuff repairs,”    Orthop. Clin. North Am. 32:501.-   78. Gartsman, G M, (1997) “Massive, irreparable tears of the rotator    cuff. Results of operative debridement and subacromial    decompression,” J. Bone Joint Surg. Am. 79:715.-   79. Gazielly, D F, Gleyze, P & Montagnon, C (1994) “Functional and    anatomical results after rotator cuff repair,” CHn. Orthop. Relat.    Res, (304):43-53.-   80. Gerber, C, Schneeberger, A G, Perren, S M & Nyffeler, R    W, (1999) “Experimental rotator cuff repair. A preliminary    study,” J. Bone Joint Surg. Am. 81:1281.-   81. Glass-Brudzinski, J, Perizzolo, D & Brunette, D M (2000)    “Effects of substratum surface topography on the organization of    cells and collagen fibers in collagen gel cultures,” J. Biomed.    Mater. Res. 61:608.-   82. Gotlin, R S & Huie, G, (2000) “Anterior cruciate ligament    injuries. Operative and rehabilitative options,” Phys. Med. Rehabil.    Clin. N. Am. 11:895-928.-   83. Goulet, F, et al., (2000) “Principles of Tissue Engineering,”    Lanza, R P, Langer, R & Vacanti, J P (eds.), pp. 639-645 Academic    Press.-   84. Grana, W A, Egle, D M, Mahnken, R & Goodhart, C W, (1994) “An    analysis of autograft fixation after anterior cruciate ligament    reconstruction in a rabbit model,” Am. J. Sports Med. 22:344-351.-   85. Gregoire, M, Orly, I, Kerebel, L M & Kerebel, B, (1987) “In    vitro effects of calcium phosphate biomaterials on fibroblastic cell    behavior,” Biol. Cell, 59:255-260.-   86. Harner, C D, et al., (1999) “Quantitative analysis of human    cruciate ligament insertions,” Arthroscopy 15:741-749 (1999).-   87. Hench, L L, (1991) “Bioceramics: from concept to clinic,” J. Am.    Cera. Soc. 74(7):1487-1510.-   88. Hynes, R O (1992) “Integrins: versatility, modulation, and    signaling in cell adhesion,” Cell, 69:11.-   89. Hynes, R O (2002) “Integrins: bidirectional, allosteric    signaling machines,” Cell, 110:673.-   90. Itoi, E, Berglund, L J, Grabowski, J J, Schultz, F M, Growney, E    S, Morrey, B F & An, K N (1995) “Tensile properties of the    supraspinatus tendon,” J. Orthop. Res. 13:578.-   91. Jackson, D W, (1992) American Academy of Orthopaedic Surgeon    Bulletin, 40:10-11.-   92. Jackson, D W, et al., (1991) Trans. Orhtop. Res. Soc. 16:208    (Abstract).-   93. Jackson, D W, et al., (1993) “A comparison of patellar tendon    autograft and allograft used for anterior cruciate ligament    reconstruction in the goat model,” Am. J. Sports Med. 21:176-185.-   94. Jackson, D W, Grood, E S, Arnoczky, S P, Butler, D L & Simon, T    M, (1987) “Cruciate reconstruction using freeze dried anterior    cruciate ligament allograft and a ligament augmentation device    (LAD). An experimental study in a goat model,” Am. J. Sports Med.    15:528-538.-   95. Jiang, J, Nicoll, S B & Lu, H H, (2003) “Effects of Osteoblast    and Chondrocyte Co-Culture on Chondrogenic and Osteoblastic    Phenotype In vitro,” Trans. Orhtop. Res. Soc. 49 (Abstract).-   96. Johnson, R J, (1982) “The anterior cruciate: a dilemma in sports    medicine,” Int. J. Sports Med. 3:71-79.-   97. Joshi, M D, Suh, J K, Marui, T & Woo, S L (1995) “Interspecies    variation of compressive biomechanical properties of the    meniscus,” J. Biomed. Mater. Res. 29:823.-   98. Karsenty, G. (1965) “A formaldehyde-glutaraldehyde fixative of    high osmolality for use in electron microscopy,” J. Cell. Biol.    27:137A.-   99. Kawamura S et al., (2005) ORS.-   100. Kim et al., (2003) Biomaterials.-   101. Knowles J C, et al., (1992) Biomaterials, 13(8):491-496.-   102. Koike, et al., (2006) “Delay of supraspinatus repair by up to    12 weeks does not impair enthesis formation: A quantitative    histologic study in rabbits.” Journal of Orthopaedic Research.    24(2):202-210.-   103. Koike, Y, Trudel, G, Uhthoff, H K, (2005) “Formation of a new    enthesis after attachment of the supraspinatus tendon: A    quantitative histologic study in rabbits,” J. Orthop. Res.    23(6):1433-1440.-   104. Kumagai, J, Sarkar, K, Uhthoff, H K, Okawara, Y, Ooshima,    A, (1994) “lmmunohistochemical distribution of type I, Il and III    collagens in the rabbit supraspinatus tendon insertion,” J. Anat.    185(Pt. 2):279-284.-   105. Kurosaka, M, Yoshiya, S & Andrish, J T, (1987) “A biomechanical    comparison of different surgical techniques of graft fixation in    anterior cruciate ligament reconstruction,” Am. J Sports Med.    15:225-229.-   106. Kurzweil, P R, Frogameni, A D & Jackson, D W, (1995) “Tibial    interference screw removal following anterior cruciate ligament    reconstruction,” Arthroscopy 11L:289-291.-   107. Kwan, M K, Wayne, J S, Woo, S L, Field, F P, Hoover, J &    Meyers, M, (1989) “Histological and biomechanical assessment of    articular cartilage from stored osteochondral shell allografts,” J.    Orthop. Res. 7:637.-   108. Langley, S M, Chai, P J, Miller, S E, Mault, J R, Jaggers, J J,    Tsui, S S, Lodge, A J, Lefurgey, A & Ungerleider, R M. (1999)    “Intermittent perfusion protects the brain during deep hypothermic    circulatory arrest,” Ann. Thorac. Surg. 68:4.-   109. Lannotti, J P, Codsi, M J, Kwon, Y W, Derwin, K, Ciccone, J &    Brems, J J (2006) “Porcine small intestine submucosa augmentation of    surgical repair of chronic two-tendon rotator cuff tears. A    randomized, controlled trial,” J. Bone Joint Surg. Am. 88:1238.-   110. Larson, R P, (1994) “The Crucial Ligaments: Diagnosis and    Treatment of Ligamentous Injuries About the Knee,” John, A. Jr. &    Feagin, J. A. (eds.), pp. 785-796 Churchill Livingstone, New York.-   111. Latridis, J C, Setton, L A, Foster, R J, Rawlins, B A,    Weidenbaum, M & Mow, V C (1998) “Degeneration affects the    anisotropic and nonlinear behaviors of human anulus fibrosus in    compression,” J. Biomech. 31:535.-   112. Lee, C H, Shin, H J, Cho, I H, Kang, Y M, Kim, I A, Park, K D &    Shin, J W (2005) “Nanofiber alignment and direction of mechanical    strain affect the ECM production of human ACL fibroblast,”    Biomaterials 26:1261.-   113. LeRoux, M A and Setton, L A (2002) “Experimental and biphasic    FEM determinations of the material properties and hydraulic    permeability of the meniscus in tension,” J. Biomech. Eng. 124:315.-   114. Li, W J, Danielson, K G, Alexander, P G & Tuan, R S “Biological    response of chondrocytes cultured in three-dimensional nanofibrous    poly(epsilon-caprolactone) scaffolds,” J. Biomed. Mater. Res.    A67:1105.-   115. Li, W J, Laurencin, C T, Caterson, E J, Tuan, R S & Ko, F    K (2002) “Electrospun nanofibrous structure: a novel scaffold for    tissue engineering,” J. Biomed. Mater. Res. 60:613.-   116. Li, W J, Mauck, R L, Cooper, J A, Yuan, X & Tuan, R S (2007)    “Engineering controllable anisotropy in electrospun biodegradable    nanofibrous scaffolds for musculoskeletal tissue engineering,” J.    Biomech. 40:1686.-   117. Li, W J, Tuli, R, Okafor, C, Derfoul, A, Danielson, K G, Hall,    D J & Tuan, R S (2005) “A three-dimensional nanofibrous scaffold for    cartilage tissue engineering using human mesenchymal stem cells,”    Biomaterials 26:599.-   118. Liu, S H, et al., (1997) “Morphology and matrix composition    during early tendon to bone healing,” Clinical Orthopaedics &    Related. Research. 253-260.-   119. Loeser, R F, Sadiev, S, Tan, L & Goldring, M B (2000) “lntegrin    expression by primary and immortalized human chondrocytes: evidence    of a differential role for α1β1 and α2β1 integrins in mediating    chondrocyte adhesion to types Il and Vl collagen,” Osteoarthritis.    Cartilage. 8:96.-   120. Loh, J C, et al., (2003) “Knee stability and graft function    following anterior cruciate ligament reconstruction: Comparison    between 11 o'clock and 10 o'clock femoral tunnel placement,”    Arthroscopy, 19:297-304.-   121. Lu, H H, Cooper, J A, Jr., Manuel, S, Freeman, J W, Attawia, M    A, Ko, F K & Laurencin, C T (2005) “Anterior cruciate ligament    regeneration using braided biodegradable scaffolds: in vitro    optimization studies,” Biomaterials 26:4805.-   122. Lu, H H, Cooper, J A, Ko, F A, Attawia, M A & Laurencin, C T,    (2001)“Effect of polymer scaffold composition on the morphology and    growth of anterior cruciate ligaments cells,” Society for    Biomaterials Proceedings (Abstract).-   123. Lu, H H, El Amin, S F, Scott, K D & Laurencin, C T, “Three    dimensional, bioactive, biodegradable, polymer bioactive glass    composite scaffolds with improved mechanical properties support    collagen synthesis and mineralization of human osteoblast-like cells    in vitro,” J. Biomed. Matls. Res. 64A:465-474.-   124. Lu, H H, et al., (2003) “Evaluation of Optimal Parameters in    the Co-Culture of Human Anterior Cruciate Ligament Fibroblasts and    Osteoblasts for Interface Tissue Engineering,” ASME 2003 Summer    Bioengineering Conference (Abstract).-   125. Lu, H H, Pollack, S R & Ducheyne, P, “45S5 bioactive glass    surface charge variations and the formation of a surface calcium    phosphate layer in a solution containing fibronectin,” J. Biomed.    Matls. Res. 54:454-461.-   126. Lu, H H, Pollack, S R & Ducheyne, P, “Temporal zeta potential    variations of 45S5 bioactive glass immersed in an electrolyte    solution,” J. Biomed. Matls. Res. 51:80-87.-   127. Ma, Z, Kotaki, M, Inai, R & Ramakrishna, S (2005) “Potential of    nanofiber matrix as tissue engineering scaffolds” Tissue Eng.    11:101.-   128. Mansat, P, Cofield, R H, Kersten, T E & Rowland, C M (1997)    “Complications of rotator cuff repair. Orthop,” Clin. North Am.    28:205.-   129. Markolf, K L, et al., (2002) “Effects of femoral tunnel    placement on knee laxity and forces in an anterior cruciate ligament    graft,” J. Orthop. Res. 20:1016-1024.-   130. Matthews, J A, Wnek, G E, Simpson, D G & Bowlin, G L (2002)    “Electrospinning of Collagen Nanofibers,” Biomacromolecules, 3:232.-   131. Matthews, L S, Soffer, S R, (1989) “Pitfalls in the use of    interference screws for anterior cruciate ligament reconstruction:    brief report,” Arthroscopy, 5:225-226.-   132. Matyas, J R, Anton, M G, Shrive, N G & Frank, C B,    (1995)“Stress governs tissue phenotype at the femoral insertion of    the rabbit MCL,” J. Biomech. 28:147-157.-   133. Mauney J R, et al., (2005) “In vitro and in vivo evaluation of    differentially demineralized cancellous bone scaffolds combined with    human bone marrow stromal cells for tissue engineering” Biomaterials    26(16):3173-3185.-   134. Mazzocca, A D, Millett, P J, Guanche, C A, Santangelo, S A &    Arciero, R A (2005) “Arthroscopic single-row versus double-row    suture anchor rotator cuff repair.” Am. J. Sports Med. 33:1861.-   135. McCarthy, D M, Tolin, B S, Schwendeman, L, Friedman, M J & Woo,    S L, (1993) “The Anterior Cruciate Ligament: Current and Future    Concepts,” Douglas, W. & M D Jackson (eds.) Raven Press, New York.-   136. Meinel L, et al., (2004) Biotechnol Bioeng, 88:379-391.-   137. Messner, K, (1997) “Postnatal development of the cruciate    ligament insertions in the rat knee. Morphological evaluation and    immunohistochemical study of collagens types I and II,” Acta    Anatomica, 160:261-268.-   138. Moffat K L, et al., (2005) ISTL-V.-   139. Moffat, et al. (2008) “Characterization of the    structure-function relationship at the ligament-to-bone interface.”    PNAS. 105(23):7947-7952.-   140. Moffat, K L, Sun, W S, Chahine, N O, Pena, P E, Doty, S B,    Hung, C T, Ateshian, G A, Lu, H H, (2006) “Characterization of the    Mechanical Properties and Mineral Distribution of the Anterior    Cruciate Ligament-to-Bone Insertion Site,” Conf. Proc. IEEE Eng.    Med. Biol. Soc. 1:2366-2369.-   141. Moore, P B & Dedman, J R., (1982) “Calcium binding proteins and    cellular regulation,” Life Sci. 31:2937-2946.-   142. Moreau J E, et al., (2005) J Orthop Res, 23:164-174.-   143. Murugan, R, & Ramakrishna, S (2007) “Design strategies of    tissue engineering scaffolds with controlled fiber orientation,”    Tissue Eng. 13:1845.-   144. Nerurkar, N L, Elliott, D M & Mauck, R L, (2007) “Mechanics of    oriented electrospun nanofibrous scaffolds for annulus fibrosus    tissue engineering,” J. Orthop. Res. 25:1018.-   145. Nicoll, S B, Wedrychowska, A, Smith, N R & Bhatnagar, R    S, (2001) “Modulation of proteoglycan and collagen profiles in human    dermal fibroblasts by high density micromass culture and treatment    with lactic acid suggests change to a chondrogenic phenotype,”    Connect. Tissue Res. 42:59-69.-   146. Niyibizi, C, Sagarrigo, V C, Gibson, G & Kavalkovich, K, (1996)    “Identification and immunolocalization of type X collagen at the    ligament-bone interface,” Biochem. Biophys. Res. Commun.    222:584-589.-   147. Noyes, F R & Barber-Westin, S D, (1996) “Revision anterior    cruciate ligament surgery: experience from Cincinnati,” Clin.    Orthop. 116-129).-   148. Noyes, F R, Mangine, R E & Barber, S, (1987) “Early knee motion    after open and arthroscopic anterior cruciate ligament    reconstruction,” Am. J. Sports Med. 15:149-160.-   149. Noyes, F R, Mangine, R E & Barber, S, (1987) “Early knee motion    after open and arthroscopic anterior cruciate ligament    reconstruction,” Am. J. Sports Med. 15:149-160.-   150. O'Brien, F J, Harley, B A, Waller, M A, Yannas, I V, Gibson, L    J & Prendergast, P J (2007) “The effect of pore size on permeability    and cell attachment in collagen scaffolds for tissue engineering,”    Technol. Health Care, 15:3.-   151. Ozaki, J, Fujimoto, S, Masuhara, K, Tamai, S & Yoshimoto,    S (1986) “Reconstruction of chronic massive rotator cuff tears with    synthetic materials,” CHn. Orthop. Relat. Res. (202):173-83.-   152. Panni, A S, Milano, G, Lucania, L & Fabbriciani, C, (1997)    “Graft healing after anterior cruciate ligament reconstruction in    rabbits,” Clin. Orthop. 203-212.-   153. Park, M C, Cadet, E R, Levine, W N, Bigliani, L U & Ahmad, C    S (2005) “Tendon-to-bone pressure distributions at a repaired    rotator cuff footprint using transosseous suture and suture anchor    fixation techniques,” Am. J. Sports Med. 33:1154.-   154. Park, M C, Tibone, J E, EIAttrache, N S, Ahmad, C S, Jun, B J &    Lee, T Q, (2007) “Part II: Biomechanical assessment for a    footprint-restoring transosseous-equivalent rotator cuff repair    technique compared with a double-row repair technique,” J. Shoulder    Elbow Surg. 16:469.-   155. Pena, F, Grontvedt, T, Brown, G A, Aune, A K & Engebretsen,    L, (1996) “Comparison of failure strength between metallic and    absorbable interference screws. Influence of insertion torque,    tunnel-bone block gap, bone mineral density, and interference,”    Am. J. Sports Med. 24:329-334.-   156. Petersen, W & Tillmann, B, (1999) “Structure and    vascularization of the cruciate ligaments of the human knee joint,”    Anat. Embryol. (Berl). 200:325-334.-   157. Pham, Q P, Sharma, U & Mikos, A G (2006) “Electrospinning of    polymeric nanofibers for tissue engineering applications: a review,”    Tissue Eng. 12:1197.-   158. Pham, Q P, Sharma, U & Mikos, A G, (2006) “Electrospun    poly(epsilon-caprolactone) microfiber and multilayer    nanofiber/microfiber scaffolds: characterization of scaffolds and    measurement of cellular infiltration,” Biomacromolecules, 7:2796.-   159. Pittenger M F, et al., (1999) Science, 284:143-147.-   160. Post, M, (1985) “Rotator cuff repair with carbon filament. A    preliminary report of five cases,” Clin. Orthop. Relat. Res.    (196):154-158.-   161. Radhakrishnan, P, Lewis, N T, Mao, J J, (2004) “Zone-specific    micromechanical properties of the extracellular matrices of growth    plate cartilage,” Ann. Biomed. Eng. 32(2):284-291.-   162. Reneker, D H and Chun, I, (1996) “Nanometer diameter fibres of    polymer, produced by electrospinning,” Nanotechnology, 7:216.-   163. Robertson, D B, Daniel, D M & Biden, E, (1986) “Soft tissue    fixation to bone,” Am. J. Sports Med. 14:398-403.-   164. Rodeo, S A, Arnoczky, S P., Torzilli, P A, Hidaka, C, Warren, R    F, (1993) “Tendon-healing in a bone tunnel. A biomechanical and    histological study in the dog.” J Bone Joint Surg Am.    75(12):1795-1803.-   165. Rodeo, S A, et al., (2006) ORS, 2006.-   166. Rodeo, S A, Suzuki, K, Deng, X H, Wozney, J & Warren, R    F, (1999) “Use of recombinant human bone morphogenetic protein-2 to    enhance tendon healing in a bone tunnel,” Am. J. Sports Med.    27:476-488.-   167. Rokito, A S, Zuckerman, J D, Gallagher, M A & Cuomo, F, (1996)    “Strength after surgical repair of the rotator cuff,” J. Shoulder    Elbow Surg. 5:12.-   168. Romeo, A A, Hang, D W, Bach, B R, Jr. & Shott, S (1999) “Repair    of full thickness rotator cuff tears. Gender, age, and other factors    affecting outcome,” CHn. Orthop. Relat. Res. (367):243-55.-   169. Safran, M R, & Harner, C D, (1996) “Technical considerations of    revision anterior cruciate ligament surgery,” Clin. Orthop. 50-64.-   170. Sagarriga, V C, Kavalkovich, K, Wu, J & Niyibizi, C, (1996)    “Biochemical analysis of collagens at the ligament-bone interface    reveals presence of cartilage-specific collagens,” Arch. Biochem.    Biophys. 328:135-142.-   171. Sahoo, S, Ouyang, H, Goh, J C, Tay, T E & Toh, S L, (2006)    “Characterization of a novel polymeric scaffold for potential    application in tendon/ligament tissue engineering.” Tissue Eng.    12:91.-   172. Sano, et al. (2002) “Experimental fascial autografting for the    supraspinatus tendon defect: Remodeling process of the grafted    fascia and the insertion into bone.” Journal of Shoulder and Elbow    Surgery. 11(2):166-173.-   173. Scapinelli, R & Little, K, (1970) “Observations on the    mechanically induced differentiation of cartilage from fibrous    connective tissue,” J. Pathol. 101:85-91.-   174. Schafer, et al., (2000) “In vitro generation of osteochondral    composites,” Biomaterials, 21:2599-2606.-   175. Schafer, et al., (2002) “Tissue-engineered composites for the    repair of large osteochondral defects,” Arthritis Rheum.    46:2524-2534.-   176. Schlegel, T F, Hawkins, R J, Lewis, C W, Motta, T & Turner, A    S, (2006) “The effects of augmentation with Swine small intestine    submucosa on tendon healing under tension: histologic and mechanical    evaluations in sheep.” Am. J. Sports Med. 34:275.-   177. Sclamberg, S G, Tibone, J E, Itamura, J M & Kasraeian,    S, (2004) “Six-month magnetic resonance imaging follow-up of large    and massive rotator cuff repairs reinforced with porcine small    intestinal submucosa.” J. Shoulder Elbow Surg. 13:538.-   178. Shellock, F G, Mink, J H., Curtin, S & Friedman, M J, (1992)    “MR imaging and metallic implants for anterior cruciate ligament    reconstruction: assessment of ferromagnetism and artifact,” J. Magn.    Reson. Imaging, 2:225-228.-   179. Shin et al., (2001) “Experimental characterization of    electrospinning: the electrically forced jet and instabilities,”    Polymer, 42(25):09955-09967.-   180. Singhvi, R, Kumar, A, Lopez, G P, Stephanopoulos, G N, Wang, D    I, Whitesides, G M & Ingber, D E, (1994) “Engineering cell shape and    function.” Science, 264:696.-   181. Sittinger, et al., (1994) “Engineering of cartilage tissue    using bieresorbable polymer carriers in perusion culture,”    Biomaterials, 15(6):451-456.-   182. Soslowsky, L J, Thomopoulos, S, Tun, S, Flanagan, C L, Keefer,    C C, Mastaw, J & Carpenter, J E, (2000) “Neer Award 1999. Overuse    activity injures the supraspinatus tendon in an animal model: a    histologic and biomechanical study.” J. Shoulder Elbow Surg. 9:79.-   183. Spalazzi J P, et al., (2006) “Development of Controlled Matrix    Heterogeneity on a Triphasic Scaffold for Orthopedic Interface    Tissue Engineering” Tissue Engineering, 12(12):3497-3508, 2006.-   184. Spalazzi J P, et al., (2006) “Elastographic Imaging of Strain    Distribution In The Anterior Cruciate Ligament and at the    Ligament-Bone Insersions” J Orthop Res, 24(10):2001-2010.-   185. Spalazzi, J P, Dionisio, K L, Jiang, J & Lu, H H, (2003)    “Chondrocyte and Osteoblast Interaction on a Degradable Polymer    Ceramic Scaffold,” ASME 2003 Summer Bioengineering Conference    (Abstract)(2003).-   186. Spalazzi, J P, Vyner, M C, Jacobs, M T, Moffat, K L, Rich, C G    & Lu, H H, (2008) “Mechanoactive Scaffold Induces Tendon Remodeling    and Expression of Fibrocartilage Markers.” CHn. Orthop. Relat. Res.    466(8):1938-1948.-   187. Stankus, J J, Guan, J, Fujimoto, K & Wagner, W R, (2006)    “Microintegrating smooth muscle cells into a biodegradable,    elastomeric fiber matrix.” Biomaterials, 27:735.-   188. Steiner, M E, Hecker, A T, Brown, C H, Jr. & Hayes, W C,    (1994)“Anterior cruciate ligament graft fixation. Comparison of    hamstring and patellar tendon grafts,” Am. J. Sports Med.    22:240-246.-   189. Sui, G, Yang, X, Mei, F, Hu, X, Chen, G, Deng, X, & Ryu,    S, (2007) “Poly-L-lactic acid/hydroxyapatite hybrid membrane for    bone tissue regeneration,” J. Biomed Mater Res A. 82(2):445-54.-   190. Teixeira, A I, Abrams, G A, Bertics, P J, Murphy, C J & Nealey,    P F, (2003) “Epithelial contact guidance on well-defined micro- and    nanostructured substrates.” J. Cell. Sci. 116:1881.-   191. Tempelhof, S, Rupp, S & Seil, R, (1999) “Age-related prevalence    of rotator cuff tears in asymptomatic shoulders.” J. Shoulder Elbow    Surg. 8:296.-   192. Thomas, N P, Turner, I G & Jones, C B, (1987) “Prosthetic    anterior cruciate ligaments in the rabbit. A comparison of four    types of replacement,” J. Bone Joint Surg. Br. 69:312-316.-   193. Thomopoulos, S, et al., (2002) “The localized expression of    extracellular matrix components in healing tendon insertion sites:    an in situ hybridization study,” J. Orthop. Res. 20:454-463.-   194. Thomopoulos, S, Fomovsky, G M, Chandran, P L & Holmes, J    W, (2007) “Collagen fiber alignment does not explain mechanical    anisotropy in fibroblast populated collagen gels.” J. Biomech. Eng.    129:642.-   195. Thomopoulos, S, Marquez, J P, Weinberger, B, Birman, V & Genin,    G M, (2006) “Collagen fiber orientation at the tendon to bone    insertion and its influence on stress concentrations.” J. Biomech.    39:1842.-   196. Thomopoulos, S, Williams, G R, Gimbel J A, Favata, M &    Soslowsky, L J, (2003) “Variations of biomechanical, structural, and    compositional properties along the tendon to bone insertion    site.” J. Orthop. Res. 21:413.-   197. Vitale, M A, Vitale, M G, Zivin, J G, Braman, J P, Bigliani, L    U & Flatow, E L, (2007) “Rotator cuff repair: an analysis of utility    scores and cost-effectiveness. J. Shoulder.” Elbow. Surg. 16:181.-   198. Wang I N E and Lu H H, ORS, 2005.-   199. Wang I N E, et al., (2006) “Age-dependent changes in matrix    composition and organization at the ligament-to-bone insertion,” J    Orthop Res, 24(8):1745-1755.-   200. Wang, I N E, Shan, J, Choi, R, Oh, S, Kepler, C K, Chen, F H,    Lu, H H, (2007) “Role of osteoblast-fibroblast interactions in the    formation of the ligament-to-bone interface.” J Orthop Res,    25(12):1609-1620.-   201. Wang, J H, Jia, F, Gilbert, T W & Woo, S L, (2003) “Cell    orientation determines the alignment of cell-produced collagenous    matrix.” J. Biomech. 36:97.-   202. Wei, X & Messner, K, (1996) “The postnatal development of the    insertions of the medial collateral ligament in the rat knee,” Anat.    Embryol. (Berl) 193:53-59.-   203. Weiler, A, Hoffmann, R F, Bail, H J, Rehm, O & Sudkamp, N    P, (2002) “Tendon healing in a bone tunnel. Part II: Histologic    analysis after biodegradable interference fit fixation in a model of    anterior cruciate ligament reconstruction in sheep,” Arthroscopy,    18:124-135.-   204. Weiler, A, Windhagen, H J, Raschke, M J, Laumeyer, A &    Hoffmann, R F, (1998) “Biodegradable interference screw fixation    exhibits pull-out force and stiffness similar to titanium screws,”    Am. J. Sports Med. 26:119-126.-   205. Weiss, J A and Maakestad, B J, (2006) “Permeability of human    medial collateral ligament in compression transverse to the collagen    fiber direction.” J. Biomech. 39:276.-   206. Williams, G R, Jr., Rockwood, C A, Jr., Bigliani, L U,    lannotti, J P & Stanwood, W, (2004) “Rotator cuff tears: why do we    repair them?” J. Bone Joint Surg. Am. 86-A:2764.-   207. Woo, S L, Gomez, M A, Seguchi, Y, Endo, C M & Akeson, W    H, (1983) “Measurement of mechanical properties of ligament    substance from a bone-ligament-bone preparation,” J. Orthop. Res.    1:22-29.-   208. Woo, S L, Maynard, J, Butler, D L, Lyon, R M, Torzilli, P A,    Akeson, W H, Cooper, R R, Oakes, B, (1988) “Ligament, Tendon, and    Joint Capsule Insertions to Bone. In Injury and Repair of the    Musculosketal Soft Tissues.” Woo, S L, Bulkwater, J A, eds.,    American Academy of Orthopaedic Surgeons: Savannah, Ga., pp.    133-166.-   209. Woo, S L, Newton, P O, MacKenna, D A & Lyon, R M, (1992) “A    comparative evaluation of the mechanical properties of the rabbit    medial collateral and anterior cruciate ligaments,” J. Biomech.    25:377-386.-   210. Wu H, et al., (2003) “Fabrication of Complex Three-Dimensional    Microchannel Systems in PDMS,” J. Am. Chem. Soc. Jan. 15, 2003;    125(2):554-559.-   211. Wuthier, R E, (1993) “Involvement of cellular metabolism of    calcium and phosphate in calcification of avian growth plate    cartilage,” J. Nutr. 123:301-309.-   212. Wutticharoenmongkol, P, Sanchavanakit, N, Pavasant, P, &    Supaphol, P (2006) “Novel bone scaffolds of electrospun    polycaprolactone fibers filled with nanoparticles. J. Nanosci    Nanotechnol.” 6(2):515-22.-   213. Xu, et al., (2004) “Aligned biodegradable nanofibrous    structure: a potential scaffold for blood vessel engineering”    Biomaterials, 25(5):877-886.-   214. Yahia, L, (1997) “Ligaments and Ligamentoplasties,” Springer    Verlag, Berlin Heidelberg.-   215. Yamanaka, K and Matsumoto, T (1994) “The joint side tear of the    rotator cuff. A followup study by arthrography.” CHn. Orthop. Relat.    Res. 304:68-73.-   216. Yang, F, Murugan, R, Wang, S & Ramakrishna, S, (2005)    “Electrospinning of nano/micro scale poly(L-lactic acid) aligned    fibers and their potential in neural tissue engineering.”    Biomaterials, 26:2603.-   217. Yin, L & Elliott, D M (2004) “A biphasic and transversely    isotropic mechanical model for tendon: application to mouse tail    fascicles in uniaxial tension.” J. Biomech. 37:907.-   218. Yoshimoto, H, Shin, Y M, Terai, H, Vacanti, J P, (2003) “A    biodegradable nanofiber scaffold by electrospinning and its    potential for bone tissue engineering.” Biomaterials,    24(12):2077-2082.-   219. Yoshiya, S, Nagano, M, Kurosaka, M, Muratsu, H & Mizuno,    K, (2000) “Graft healing in the bone tunnel in anterior cruciate    ligament reconstruction,” Clin. Orthop. 278-286.-   220. Zhong, S, Teo, W E, Zhu, X, Beuerman, R W, Ramakrishna, S &    Yung, L Y, (20060 “An aligned nanofibrous collagen scaffold by    electrospinning and its effects on in vitro fibroblast culture.” J.    Biomed. Mater. Res. A79:456.

What is claimed is:
 1. An implantable device for soft-tissue or softtissue-to-bone fixation, repair, augmentation, or replacement comprisinga biomimetic scaffold, said scaffold being biphasic and continuous froma first phase to a second phase, wherein fibers in the first phase haveat least one of different alignment, different orientation, differentcomposition and different coating as compared to fibers in the secondphase, and wherein the fibers in the first phase are made from a polymerand the second phase is coupled to the first phase, the fibers in thesecond phase are composite fibers, each of the composite fibers of thesecond phase is made from a combination of a polymer and a biocompatibleceramic.
 2. The implantable device according to claim 1, wherein thebiocompatible ceramic is selected from the group consisting of siliconnitride-based ceramics, Pseudowollastonite ceramics (β-CaSiO₃),bredigite (Ca₇MgSi₄O₁₆) ceramics, monophase ceramics of monticellite(CaMgSiO(4)), akermanite ceramics (Ca₂MgSi₂O₇), tricalcium silicate(Ca(3)SiO(5)), hydroxyapatite, bio-active glass, calcium phosphate,dense calcium sulfate (DCaS), porous silicated calcium phosphate(Si—CaP), tricalcium phosphate (TCP), calcium pyrophosphate (CPP), andcombinations thereof.
 3. The implantable device according to claim 1,wherein at least one of the phases further comprises a bioactive agentselected from the group consisting of an anti-infective, anextracellular matrix component, an antibiotic, bisphosphonate, ahormone, an analgesic, an anti-inflammatory agent, a growth factor, anangiogenic factor, a chemotherapeutic agent, an anti-rejection agent, anRGD peptide, and combinations thereof.
 4. The implantable deviceaccording to claim 3, wherein the growth factor is selected from thegroup consisting of a member of the Transforming Growth Factor (TGF)super family, a vascular endothelial growth factor (VEGF), aplatelet-derived growth factor (PDGF), an insulin-derived growth factor(IGF), a modulator of a growth factor, and combinations thereof.
 5. Theimplantable device according to claim 4, wherein a member of the TGFsuper family is selected from the group consisting of TGF-β, bonemorphogenetic proteins (BMPs), growth differentiation factors (GDFs),Activin A and Activin B, lnhibin A, lnhibin B, anti-mullerian hormone,Nodal, and combinations thereof.
 6. The implantable device according toclaim 5, wherein the TGF-β is selected from the group consisting ofTGF-β1, TGF-β2, TGF-β3, and combinations thereof.
 7. The implantabledevice according to claim 5, wherein the BMP is selected from the groupconsisting of BMP 1-20 and combinations thereof.
 8. The implantabledevice according to claim 5, wherein the GDFs are selected from thegroup consisting of GDF1-15 and combinations thereof.
 9. The implantabledevice according to claim 4, wherein the IGF is selected from the groupconsisting of IGF1, IGF2, insulin growth factor binding proteins 1-6(IGFBP1-6), and combinations thereof.
 10. The implantable deviceaccording to claim 4, wherein a modulator of a growth factor is a SMAD(small mothers against decapentaplegic) selected from the groupconsisting of SMAD1-9 and combinations thereof.
 11. The implantabledevice according to claim 1, further comprising a hydrogel disposed onat least a portion of one or both of the phases.
 12. The implantabledevice according to claim 11, wherein the hydrogel is composed of amaterial selected from the group consisting of agarose, carrageenan,polyethylene oxide, polyethylene glycol, tetraethylene glycol,triethylene glycol, trimethylolpropane ethoxylate, pentaerythritolethoxylate, hyaluronic acid, thiosulfonate polymer derivatives,polyvinylpyrrolidone-polyethylene glycol-agar, collagen, dextran,heparin, hydroxyalkyl cellulose, chondroitin sulfate, dermatan sulfate,heparan sulfate, keratan sulfate, dextran sulfate, pentosan polysulfate,chitosan, alginates, pectins, agars, glucomannans, galactomannans,maltodextrin, amylose, polyalditol, alginate-based gels cross-linkedwith calcium, polymeric chains of methoxypoly(ethylene glycol)monomethacrylate, chitin, poly(hydroxyalkyl methacrylate),poly(electrolyte complexes), poly(vinylacetate) cross-linked withhydrolysable bonds, water-swellable N-vinyl lactams, carbomer resins,starch graft copolymers, acrylate polymers, polyacrylamides, polyacrylicacid, ester cross-linked polyglucans, and derivatives and combinationsthereof.
 13. The implantable device according to claim 11 furthercomprising fibroblasts, chondrocytes, osteoblasts, osteoblast-likecells, stem cells, and combinations thereof.
 14. The implantable deviceaccording to claim 13, wherein fibroblasts, chondrocytes, stem cells,and combinations thereof are disposed on at least a portion of the firstphase.
 15. The implantable device according to claim 13, whereinchondrocytes, osteoblasts, osteoblast-like cells, stem cells, andcombinations thereof are disposed on at least a portion of the secondphase.
 16. The implantable device according to claim 13, whereinfibroblasts, stem cells, and chondrocytes are disposed on at least aportion of the first phase and chondrocytes, osteoblasts,osteoblast-like cells, stem cells, and combinations thereof are disposedon at least a portion of the second phase.
 17. The implantable deviceaccording to claim 13, wherein the stem cells are undifferentiated. 18.The implantable device according to claim 13, wherein the stem cells arepre-differentiated prior to disposition on the implantable device. 19.The implantable device according to claim 1, wherein the fibers in thefirst phase include nanofibers and the fibers in the second phaseinclude nanofibers.
 20. The implantable device according to claim 1,wherein the second phase further comprises particles or nanoparticlesmade from a biocompatible ceramic.
 21. An implantable biphasicbiomimetic device for soft-tissue or soft tissue-to-bone interfacefixation, repair, augmentation, or replacement, comprising a first phasecomprising fibers made from a polymer and a second phase coupled to thefirst phase, said second phase comprising composite fibers, and each ofthe composite fibers of the second phase is made from a combination of apolymer and a biocompatible ceramic, wherein the first and second phasesare continuous.
 22. The implantable device according to claim 21,wherein the polymer is selected from the group consisting of aliphaticpolyesters, poly(amino acids), modified proteins, polydepsipeptides,copoly(ether-esters), polyurethanes, polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(ε-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, modified polysaccharides,polycarbonates, polytyrosinecarbonates, polyorthocarbonates,poly(trimethylene carbonate), poly(phosphoester)s, polyglycolide,polylactides, polyhydroxybutyrates, polyhydroxyvalerates,polydioxanones, polyalkylene Oxalates, polyalkylene succinates,poly(malic acid), poly(maleic anhydride), polyvinylalcohol,polyesteramides, polycyanoacrylates, polyfumarates, poly(ethyleneglycol), polyoxaesters containing amine groups,poly(lactide-co-glycolides), poly(lactic acid)s, poly(glycolic acid)s,poly(dioxanone)s, poly(alkylene alkylate)s, biopolymers, collagen, silk,chitosan, alginate, and a blend of two or more of the precedingpolymers.
 23. The implantable device according to claim 21, wherein thebiocompatible ceramic is selected from the group consisting of siliconnitride-based ceramics, Pseudowollastonite ceramics (β-CaSiO₃),bredigite (Ca₇MgSi₄O₁₆) ceramics, mono-phase ceramics of monticellite(CaMgSiO(4)), akermanite ceramics (Ca₂MgSi₂O₇), tricalcium silicate(Ca(3)SiO(5)), hydroxyapatite, bio-active glass, calcium phosphate,dense calcium sulfate (DCaS), porous silicated calcium phosphate(Si—CaP), tricalcium phosphate (TCP), calcium pyrophosphate (CPP), andcombinations thereof.
 24. The implantable device according to claim 21,wherein at least one of the phases further comprises a bioactive agentselected from the group consisting of an anti-infective, anextracellular matrix component, an antibiotic, bisphosphonate, ahormone, an analgesic, an anti-inflammatory agent, a growth factor, anangiogenic factor, a chemotherapeutic agent, an anti-rejection agent, anRGD peptide, and combinations thereof.
 25. The implantable deviceaccording to claim 24, wherein the growth factor is selected from thegroup consisting of a member of the Transforming Growth Factor (TGF)super family, a vascular endothelial growth factor (VEGF), aplatelet-derived growth factor (PDGF), an insulin-derived growth factor(IGF), a modulator of a growth factor, and combinations thereof.
 26. Theimplantable device according to claim 25, wherein a member of the TGFsuper family is selected from the group consisting of TGF-β, bonemorphogenetic proteins (BMPs), growth differentiation factors (GDFs),Activin A and Activin B, lnhibin A, lnhibin B, anti-mullerian hormone,Nodal, and combinations thereof.
 27. The implantable device according toclaim 26, wherein the TGF-β is selected from the group consisting ofTGF-β1, TGF-β2, TGF-β3, and combinations thereof.
 28. The implantabledevice according to claim 26, wherein the BMP is selected from the groupconsisting of BMP1-20 and combinations thereof.
 29. The implantabledevice according to claim 26, wherein the GDFs are selected from thegroup consisting of GDF1-15 and combinations thereof.
 30. Theimplantable device according to claim 25, wherein the IGF is selectedfrom the group consisting of IGF1, IGF2, insulin growth factor bindingproteins 1-6 (IGFBP1-6), and combinations thereof.
 31. The implantabledevice according to claim 25, wherein a modulator of a growth factor isa SMAD (small mothers against decapentaplegic) selected from the groupconsisting of SMAD1-9 and combinations thereof.
 32. The implantabledevice according to claim 21, further comprising a hydrogel disposed onat least a portion of one or both of the phases.
 33. The implantabledevice according to claim 32, wherein the hydrogel is composed of amaterial selected from the group consisting of agarose, carrageenan,polyethylene oxide, polyethylene glycol, tetraethylene glycol,triethylene glycol, trimethylolpropane ethoxylate, pentaerythritolethoxylate, hyaluronic acid, thiosulfonate polymer derivatives,polyvinylpyrrolidone-polyethylene glycol-agar, collagen, dextran,heparin, hydroxyalkyl cellulose, chondroitin sulfate, dermatan sulfate,heparan sulfate, keratan sulfate, dextran sulfate, pentosan polysulfate,chitosan, alginates, pectins, agars, glucomannans, galactomannans,maltodextrin, amylose, polyalditol, alginate-based gels cross-linkedwith calcium, polymeric chains of methoxypoly(ethylene glycol)monomethacrylate, chitin, poly(hydroxyalkyl methacrylate),poly(electrolyte complexes), poly(vinylacetate) cross-linked withhydrolysable bonds, water-swellable N-vinyl lactams, carbomer resins,starch graft copolymers, acrylate polymers, polyacrylamides, polyacrylicacid, ester cross-linked polyglucans, and derivatives and combinationsthereof.
 34. The implantable device according to claim 33, furthercomprising fibroblasts, chondrocytes, osteoblasts, osteoblast-likecells, stem cells, and combinations thereof.
 35. The implantable deviceaccording to claim 34, wherein fibroblasts, stem cells, chondrocytes,and combinations thereof are disposed on at least a portion of the firstphase.
 36. The implantable device according to claim 34, whereinchondrocytes, osteoblasts, osteoblast-like cells, stem cells, andcombinations thereof are disposed on at least a portion of the secondphase.
 37. The implantable device according to claim 34, whereinfibroblasts, stem cells, and chondrocytes are disposed on at least aportion of the first phase and chondrocytes, osteoblasts,osteoblast-like cells, stem cells, and combinations thereof are disposedon at least a portion of the second phase.
 38. The implantable deviceaccording to claim 34, wherein the stem cells are undifferentiated. 39.The implantable device according to claim 34, wherein the stem cells arepre-differentiated prior to disposition on the implantable device. 40.The implantable device according to claim 21, which is about 0.2 toabout 2.0 mm thick.
 41. The implantable device according to claim 40,which is about 5.0 cm in length and about 5.0 cm in width.
 42. Theimplantable device according to claim 40, which has a diameter of about10 cm.
 43. The implantable device according to claim 21, wherein thefibers in the first phase include nanofibers and the fibers in thesecond phase include nanofibers.
 44. The implantable device according toclaim 21, wherein the second phase further comprises particles ornanoparticles made from a biocompatible ceramic.
 45. An implantabledevice for fixation, repair, augmentation, or replacement of a rotatorcuff or a tendon-to bone interface thereof comprising a biphasic andbiomimetic scaffold having a first phase comprising fibers whoseanisotropy mimics that of a tendon and non-mineralized fibrocartilage,said fibers having been made from a polymer, and a second phase coupledto the first phase, said second phase comprising composite fibers whoseanisotropy mimics that of mineralized fibrocartilage and bone, each ofsaid composite fibers is made from a combination of a polymer and abiocompatible ceramic, wherein the first and second phases arecontinuous.
 46. The implantable device according to claim 45, whereinthe fibers in the first phase include nanofibers and the fibers in thesecond phase include nanofibers.
 47. The implantable device according toclaim 45, wherein the second phase further comprises particles ornanoparticles made from a biocompatible ceramic.
 48. An implantabledevice for soft-tissue or soft tissue-to-bone fixation, repair,augmentation, or replacement comprising a biomimetic scaffold, saidscaffold comprising at least two adjacent continuous phases, wherein afirst phase includes fibers having at least one of different alignment,different orientation, different composition and different coating ascompared to fibers in a second phase, wherein the first phase comprisescomposite fibers, each composite fiber is made from a combination of apolymer and a biocompatible ceramic, the second phase is coupled to thefirst phase, said second phase comprising fibers made from a polymer,and further comprising a third phase coupled to the second phase, whichthird phase comprises composite fibers, each composite fiber is madefrom a combination of a polymer and a biocompatible ceramic, wherein thefirst, second and third phases are continuous.
 49. The implantabledevice according to claim 48, wherein the fibers in the first phaseinclude nanofibers and the fibers in the second phase includenanofibers.
 50. The implantable device according to claim 48, whereinthe first phase, the third phase or both the first and the third phasefurther comprises particles or nanoparticles made from a biocompatibleceramic.